Medical Electron Accelerators
Notice Medicine is an ever-changing science. As new research and clinical experience broaden our knowledge, changes in treatment and drug therapy are required. The authors and the publisher of this work have checked with sources believed to be reliable in their efforts to provide information that is complete and generally in accord with the standards accepted at the time of publication. However, in view of the possibility of human error or changes in medical sciences, neither the editors nor the publisher nor any other party who has been involved in the preparation or publication of this work warrants that the information contained herein is in every respect accurate or complete, and they are not responsible for any errors or omissions or for the results obtained from use of such information. Readers are encouraged to confirm the information contained herein with other sources. For example and in particular, readers are advised to check the product information sheet included in the package of each drug they plan to administer to be certain that the information contained in this book is accurate and that changes have not been made in the recommended dose or in the contraindications for administration. This recommendation is of particular importance in connection with new or infrequently used drugs.
Medical Electron Accelerators Department of Radiation Oncology Stanford University School of Medicine Stanford, California
Craig S. Nunan and Eiji Tanabe VarianAssociates Palo Alto, California
McGRAW-HILL, INC. Health Professions Divisiorz New York St. Louis Sun Francisco Auckland Bogota' Caracas Lisbor~ London Madrid Mexico Milan Montreal New Delhi Paris Sun Juan Singapore Sydney Tokyo Toronto
Medical Electron Accelerators Copyright O 1993 by McGraw-Hill, Inc. All rights reserved. Printed in the United States of America. Except as permitted under the United States Copyright Act of 1976, no part of this publication may be reproduced or distributed in any form or by any means, or stored in a data base or retrieval system, without the prior written permission of the publisher. 1234567890 HALHAL 98765432 ISBN 0-07-105410-3 This book was set in Times Roman by Northeastern Graphic Services, Inc. The editors were Jane Pennington and Steven Melvin; the production supervisor was Richard Ruzycka; the cover designer was Marsha Cohenffarallelogram. Arcata GraphicsMalliday was printer and binder.
Library of Congress Cataloging-in-Publication Data Karzmark, C. J. Medical electron accelerators 1 C.J. Karzmark, Craig S. Nunan, and Eiji Tanabe. p. cm. Includes bibliographical references and index. ISBN 0-07- 105410-3 : 1. Cancer-Radiotherapy. 2. Electron accelerators. 3. ElectronsTherapeutic use. I. Nunan, Craig S. 11. Tanabe, Eiji. III. Title. 3. Particle [DNLM: 1. Electrons. 2. Neoplasms-radiotherapy. QZ 269 K18mI Accelerators. RC271.E43K37 1993 6 16.99'40642-dc20 DNLMIDLC for Library of Congress 9249307 CIP
Contents
CHAPTER
1
The Medical Electron Accelerator OVERVIEW
1
1
THE NEED FOR MEDICAL ELECTRON ACCELERATORS 1 1 OUTLINE OF THIS BOOK GOALS OF RADIOTHERAPY 1 IMPACT OF TREATMENT COURSE FRACTIONATION ON MACHINE 2 PERFORMANCE REQUIREMENTS 3 USER PREFERENCES FOR BEAM MODE TREATMENT BEAM GENERATION 4
HISTORY OF ELECTRON ACCELERATORS
6
DIRECT ACCELERATORS 6 BETATRONS 6 MICROWAVE ELECTRON LINEAR ACCELERATORS (LINACS) RECIRCULATING ELECTRON ACCELERATORS 13
ELEMENTARY DESCRIPTION OF MEDICAL LINACS
7
15
MICROWAVE ACCELERATION PRINCIPLE 16 BEAM CURRENT REQUIREMENTS 18 IN X-RAY MODE MAJOR SUBSYSTEMS AND COMPONENTS 18 20 SUMMARY OF ENERGY CONVERSION STEPS
DESIGN CRITERIA FOR RADIOTHERAPY ACCELERATORS
20
CLINICAL REQUIREMENTS 22 SOME DESIGN CHALLENGES 24 CHANGES IN TECHNOLOGY FROM EARLY-TO-MODERN MACHINES SUMMARY: ACCELERATOR MAJOR SUBSYSTEMS 26
ONE FUTURE DIRECTION OF EQUIPMENT DEVELOPMENT IN RADIATION THERAPY 27 REFERENCES
CHAPTER
2
29
Radiotherapy Modalities
33
ORTHOVOLTAGE X-RAY THERAPY
33
26
vi
CONTENTS
MEGAVOLTAGE X-RAY THERAPY
34
TOTAL-BODY AND HEMIBODY X-RAY THERAPY (MAGNA-FIELD THERAPY) 35 MEGAVOLTAGE ELECTRON THERAPY TOTAL SKIN ELECTRON THERAPY
37
38
INTRAOPERATIVE RADIATION THERAPY ARCTHERAPY
39
41
DYNAMIC AND CONFORMAL THERAPY AND MULTILEAF COLLIMATORS 41 STEREOTACTIC RADIOSURGERY REFERENCES
CHAPTER.
3
43
45
49
Microwave Principles for Linacs ELEMENTARY LINAC MICROWAVES
50
50
TRANSMISSION LINES AND WAVEGUIDES
51
IMPEDANCE MATCHING AND VOLTAGE STANDING WAVE RATIO RESONANCE AND RESONANT CAVITIES
55
PHASE VELOCITY AND GROUP VELOCITY PERIODIC STRUCTURES AND COUPLING MODE AND DISPERSION
CHAPTER
4
59
62
SHUNT IMPEDANCE AND TRANSIT TIME REFERENCES
59
64
66
67
Microwave Accelerator Structures ELECTRON GUNS AND INJECTION 67 CATHODE 67 DESIGN OF AN ELECTRON GUN 68
ELECTRON INTERACTION WITH MICROWAVE FIELD MOTION OF ELECTRONS SPACE HARMONICS 70
68
TRAVELING-WAVE ACCELERATORS
70
THEORY OF OPERATION 70 STRUCTURES 71 ELECTRON INJECTION AND BUNCHING BEAM LOADING AND LOAD LINE 75
STANDING-WAVEACCELERATORS
72
76
THEORY OF OPERATION 76 STRUCTURES 78 ELECTRON INJECTION AND BUNCHING
79
68
54
vii
CONTENTS
BEAM LOADING AND LOAD LINE
80
TRAVELING-WAVE VERSUS STANDING-WAVE ACCELERATORS DESIGN OF ACCELERATOR CAVITIES REFERENCES
CHAPTER
5
86
87
Microwave Power Sources and Systems MAGNETRONS KLYSTRONS
82
89
89 91
RADIO FREQUENCY DRIVERS CIRCULATORS
92
95
OTHER MICROWAVE COMPONENTS
97
WAVEGUIDE BENDS AND TWISTS, AND FLEXIBLE WAVEGUIDES DIRECTIONAL COUPLERS 98 98 SHUNT, SERIES, AND HYBRID TEES ROTARY JOINTS 99 WAVEGUIDE WINDOWS 100 WATER LOADS 101
AUTOMATIC FREQUENCY CONTROL
102
LOW ENERGY (MAGNETRON) AUTOMATIC FREQUENCY CONTROL HIGH ENERGY (KLYSTRON) AUTOMATIC FREQUENCY CONTROL
REFERENCES
CHAPTER
6
97
102 103
104
Pulse Modulators and Auxiliary Systems PULSE MODULATORS VACUUM SYSTEMS
105
105 107
WATER COOLING SYSTEM MISCELLANEOUS SYSTEMS
110 111
GAS DIELECTRIC SYSTEM 111 PNEUMATIC SYSTEM 113
REFERENCES
CHAPTER
7
113
Beam Optics of Magnet Systems OVERVIEW
115
115
STRAIGHT AHEAD LINACS 115 BENT BEAM LINACS 115 LINAC BEAM CHARACTERISTICS 115 EFFECT OF MAGNET SYSTEM CHOICE ON ISOCENTER HEIGHT
ELECTRON MOTION IN MAGNETIC FIELDS
116
ELECTRON MOMENTUM 116 ELECTRON MOTION IN THE DIPOLE MAGNETIC FIELD
116
115
viii
CONTENTS
ELECTRON MOTION IN THE FRINGE FIELD AT THE EDGE OF THE DIPOLE MAGNET 117 119 ELECTRON MOTION IN THE QUADRUPOLE MAGNETIC FIELD 120 ELECTRON MOTION IN THE MAGNETIC FIELD OF A SOLENOID BEAM STEERING COILS 122 BEAM TRANSPORT 122
BEAM EMITTANCE
122
NONACHROMATIC BEND MAGNET SYSTEMS 125 ACHROMATIC BEND MAGNET SYSTEMS 129 SYMMETRICAL270" SINGLE SECTOR HYPERBOLIC POLE GAP 129 129 SYMMETRICAL270" SINGLE SECTOR LOCALLY TILTED POLE GAP 129 SYMMETRICAL 270" SINGLE SECTOR STEPPED POLE GAP SYMMETRICAL270" THREE SECTOR UNIFORM POLE GAP, TWO Cx CROSS-OVERS 130 SYMMETRICAL270" THREE SECTOR UNIFORM POLE GAP, ONE Cx CROSS-OVER 131 133 ASYMMETRIC 270" TWO SECTOR UNIFORM POLE GAP ASYMMETRIC 112 115" THREE SECTOR UNIFORM POLE GAP 133 SYMMETRICAL 180" FOUR SECTOR UNIFORM POLE GAP-ISOCHRONOUS 133 REFERENCES
CHAPTER.
8
134
Treatment Beam Production
137
GEOMETRIC RESTRICTIONS OF RADIATION HEAD ANCILLARY COMPONENTS 139 RADIATION SHIELDING 140 BEAM COLLIMATORS 141 FIELD LIGHT AND RANGEFINDER
138
142
ELECTRON THERAPY 142 ELECTRON SCATTERING SYSTEM 144 ELECTRON SCANNING SYSTEM 144 MICROTRONS VERSUS LINACS FOR ELECTRON THERAPY
X-RAYTHERAPY
146 X-RAY TARGET AND FLATTENING FILTER X-RAY SCANNING SYSTEM 148 SCANNED BEAM DOSIMETRY 150
CONTAMINATION OF RADIATION BEAM
147
150
NEUTRON LEAKAGE AND RADIOACTIVATION REFERENCES
CHAPTER.
9
150
151
Dose Monitoring and Beam Stabilization IONIZATION CHAMBER
157
157
145
ix
CONTENTS
160
INTEGRATED DOSE AND DOSE RATE FIELD UNIFORMITY CONTROL
161
MONITORING AND CONTROL OF MULTIMODALITY TREATMENT UNITS 162 TREATMENT BEAM STABILIZATION
164
ELECTRICAL AND MAGNETIC INTERFERENCE REFERENCES
CHAPTER
10
166
167
169
Accelerator Control and Safety Interlocking COMPUTER CONTROL MINIATURIZATION
170 170
SEMICONDUCTOR DEVICES AND ELECTRICAL INTERFERENCE ACCELERATOR OPERATIONAL STATES INTERLOCK SYSTEM
173
173
PROTECTION AGAINST EXTREME DOSE CONTROL CONSOLE
176
177
MOTION CONTROL SYSTEM RECORD AND VERIFY SYSTEM PATIENT RECORD KEEPING
178 180 181
COMPUTER INTEGRATION OF RADIOTHERAPY REFERENCES
CHAPTER
11
171
181
185
Multi-X-Ray Energy Accelerators DESIGN CHALLENGES
I89
189
CLINICAL NEED 189 PERFORMANCE REQUIREMENTS 189 ELECTRON BEAM DURING ACCELERATION 190 ENERGY STABILITY 190 DOSE SPATIAL DISTRIBUTION AND CALIBRATION IN INITIAL SECONDS 190
EQUIPMENT DESIGN ALTERNATIVES
191
MICROWAVE POWER SOURCE-KLY STRON VERSUS MAGNETRON 191 ELECTRON GUN-TRIODE VERSUS DIODE 191 ACCELERATOR GUIDE-TRAVELING WAVE VERSUS STANDING WAVE 191 SWITCHING FROM HIGH TO LOW ENERGY IN A TRAVELING WAVE GUIDE 193 SWITCHING FROM HIGH TO LOW X-RAY ENERGY IN A STANDING WAVE GUIDE 193 BEAM LOADING 197 NON-CONTACT TYPE SIDE CAVITY ENERGY SWITCH 197
x
CONTENTS
SYSTEM FEEDBACK CONTROL PHILOSOPHY
REFERENCES
CHAPTER .12
199
199
Patient Support Assembly and Treatment Accessories
201
PATIENT SUPPORT ASSEMBLY 201 PATIENT TABLE SUPPORT TYPES 201 TABLETOP 203 TREATMENT CHAIR 203 TREATMENT ACCESSORIES
204
FIELD SHAPING SYSTEMS 204 WEDGE FILTERS AND TISSUE COMPENSATORS 205 PATIENT CONTOUR SYSTEMS 207 PATIENT IMMOBILIZATION DEVICES 208 MECHANICAL AND OPTICAL POINTERS 208 PATIENT POSITIONING AND MOTION DETECTION 208
REFERENCES
CHAPTER .13
209
Treatment Simulators, Treatment Planning and Portal Imaging TREATMENT SIMULATORS
213
MECHANICAL FEATURES 2 14 RADIOGRAPHY AND FLUOROSCOPY 216 SIMULATION ACCESSORIES 2 17 OPERATIONAL ORGANIZATION 2 17 REGULATORY REQUIREMENTS 2 17 SIMULATOR USAGE 219 CONTEMPORARY DEVELOPMENTS 219
TREATMENT PLANNING RESOURCES
220
222
RADIOGRAPHIC (FILM) PORTAL IMAGING PHYSICS OF CONVENTIONAL PORT FILMING PORTAL FILM ENHANCEMENT TECHNIQUES
ELECTRONIC PORTAL IMAGING
224 224 226
227
VALUE OF ELECTRONIC IMAGING 227 ONE-DIMENSIONAL VERSUS TWO-DIMENSIONAL DETECTORS 228 SILICON DIODE LINEAR ARRAY-MECHANICALLY SCANNED 230 MULTIWIRE SEQUENTIALLY PULSED (ELECTRONICALLY SCANNED) LIQUID IONIZATION CHAMBER 230 MECHANICALLY ROTATED MULTICHANNEL IONIZATION CHAMBER TAPERED FIBER OPTICS TO TV CAMERA 232 233 LENS TO TV CAMERA TWO-DIMENSIONAL ARRAY OF SILICON DETECTORS 234 TWO-DIMENSIONAL AMORPHOUS SILICON ARRAY 234 MOUNTING A DETECTOR ON A LINAC 235
213
xi
CONTENTS
PHOTON SPECTRUM IN PORTAL IMAGING
235
DEPENDENCE OF IMAGE CONTRAST ON X-RAY ENERGY OFF-AXIS PORTAL X-RAY TUBE 236 ON-AXIS PORTAL X-RAY TUBE 236
REFERENCES
CHAPTER
.14
236
237
Radiotherapy Accelerator Facilities
241
FACILITY PLANNING AND OPERATIONAL RESOURCES
241
MEGAVOLTAGE THERAPY ACCELERATORS AND TREATMENT FACILITIES 244 MULTIMODALITY THERAPY INSTALLATION 244 TREATMENT ROOM DESIGN 247 SHIELDING BARRIER DESIGN 247 ENTRY DOORS AND MAZES 25 1 PATIENT OBSERVATION AND COMMUNICATION 25 1 RADIOACTIVE AND TOXIC GAS PRODUCTION 252 RADIOACTIVATION OF PATIENT 252
ACCELERATOR MAINTENANCE AND USAGE
252
CONVENTIONAL MAINTENANCE 252 EXPERT SYSTEMS 254 TEST EQUIPMENT AND INSTRUMENTATION 254 PERIODIC TESTS OF FUNCTIONAL PERFORMANCE USAGE AND DOWNTIME EXPERIENCE 256
255
SAFETY ASPECTS-FACILITY AND MACHINE INTERLOCKS HUMAN ENGINEERING ASPECTS REFERENCES
CHAPTER
.15
256
257
258
Medical Microtron Accelerators CIRCULAR MICROTRON
261
261
CAVITY POWER REQUIREMENTS 26 1 MAGNET SIZE 262 INJECTION METHODS FOR INCREASED ENERGY PER ORBIT PHASE STABILITY 262 BEAM CURRENT AND FOCUSING 263 GANTRY 263 MACHINES FOR RADIOTHERAPY 263
RACETRACK MICROTRON
263
CONFIGURATION 263 FOCUSING 264 ACCELERATOR STRUCTURE POWER INJECTION 264 EXTRACTION 265 ALIGNMENT PRECISION 265
264
262
xii
CONTENTS
MACHINE FOR RADIOTHERAPY
REFERENCES
CHAPTER
.I6
266
266
Other Types of Medical Electron Accelerators HISTORY
267
TRANSFORMER-RECTIFIERUNITS RESONANT TRANSFORMER
268
270
VAN DE GRAAFF GENERATOR
270
BETATRON AND ELECTRON SYNCHROTRON REFERENCES
APPENDIX
A
267
271
274
Generation of Radiation Beams
275
X-RAY BEAMS 275 PHOTON SPECTRA ON THE AXIS OF AN UNFLATI'ENED LOBE 275 ANGULAR DISTRIBUTION OF PHOTON INTENSITY 275 CHOICE OF TARGET MATEIUAL AND THICKNESS 276 CHOICE OF MATERIALS AND THEIR DISTRIBUTION IN THE X-RAY FLATI'ENING FILTER 278 ELECTRON BEAMS
278
SPURIOUS SOURCES 278 SCATTERING FOILS 279 EFFECT OF ENERGY SPECTRUM WIDTH ON THE SLOPE OF DEPTH DOSE CURVE 279 RADIATION INTERACTIONS IN THE PATIENT 279
ADDITIONAL MATERIAL FOR RADIATION CALCULATIONS REFERENCES
APPENDIX.
B
281
Survey of Medical Linacs
287
HISTORICAL SUMMARY OF MANUFACTURERS' TYPES CONTEMPORARY RADIOTHERAPY ACCELERATORS REFERENCES
APPENDIX
C
287
Miscellaneous
297
C-1 ABBREVIATIONS C-2 SYMBOLS
297
298
C-3 GREEK SYMBOLS C-4 UNITS
299
C-5 TERMINOLOGY
Index
299
309
300
287 287
281
Preface
The emphasis in this book is on the design and principles of operation of microwave electron linear accelerators for the treatment of cancer. Associated equipment and accessories in the radiotherapy clinic are described, such as simulators, treatmentplanning units, radiotherapy management systems, multileaf collimators, and electronic portal imagers. Other electron accelerators, including rnicrotrons, betatrons, and direct accelerators, are also discussed. The intended audience for this book includes medical physicists and engineers, radiation technologists, radiation oncologists and residents, hospital and radiology administrators, design and service engineers, as well as university students in physics and engineering. A knowledge of the various engineering tradeoffs in machine design can assist users in the initial selection of the appropriate machine for their application. An understanding of the principles of accelerator operation and their application promotes confidence in their more effective use. This book is useful directly as a teaching aid for physicists and engineers in training, medical residents, and radiation therapy technologists. The book is particularly designed for study by persons without extensive knowledge and experience of accelerator technology. It is also organized to serve as a ready reference. We have assumed on the part of the reader only a knowledge of elementary physics and mathematics. Emphasis is placed on how accelerators function and how they are used in the treatment of cancer. Illustrations, tables, and analogies are abundantly used for clarity to the nonspecialist. A solid theoretical base is provided for the specialist by descriptive text and illustration. The subject matter includes a history of development and application, general theory of acceleration, accelerator systems, radiation beam systems and associated equipment, performance characteristics, testing, and use. The major modules of a representative medical accelerator are described, including the principles of operation, and how these modules function collectively to produce electron and x-ray beams for radiotherapy. Electron accelerators were initially developed in research laboratories for use in experimental nuclear physics, and an
extensive literature exists on that early work. These physics laboratories subsequently moved on to the development of much higher energy particle accelerators for elementary particle physics. Meanwhile, manufacturers have carried on the development of these more moderate energy accelerators for pragmatic applications. The formal literature on these developments by laboratory researchers, manufacturers, and users is sparse and scattered. The goal of this book is to provide in one convenient place a lucid description of the design and operation of medical electron accelerators, together with extensive references for a more detailed study of specific areas of the readers' interest. Early medical accelerator developments were primarily concerned with the transition from a supportive physics laboratory environment to reliable operation in the hospital clinic where the laboratory resources were largely unavailable. Over a period of time advances in vacuum technology and microwave, electronic, and mechanical systems contributed significantly to the precision, reliability, and stability of these treatment machines, as well as to markedly improved treatment beam characteristics. In recent years, the number of these accelerators in use has increased rapidly. Their technology is continually advancing to meet the developing requirements of radiotherapy. Present emphasis is on features concerned with how best to treat the widest possible group of patients safely, rapidly, and comfortably and on facilitating efficient management of equipment operation in the clinic. The application of computers and imaging technology is being increasingly incorporated in this effort. The preparation of this book would not have been possible without the valuable assistance of many colleagues: engineers, physicists, physicians, and technologists in clinics, in industry, and universities. Often they have provided insight and understanding pertaining to specific considerations related to the design and application of these accelerators. We are grateful for the contributions from the following individuals at Varian Associates: David Auerbach, Steve Cheung, Verne Edwardson, Dan Hardesty, Joseph Jachinowski, Stan Johnsen, Martin Kandes, Phil LaRiviere, Dick Levy, Ray McIntyre, Stan Mansfield, David Maurier, David Penning, Niel Pering, Dick
xiv
PREFACE
Thompson, and Gene Tochilin, and especially for the technical assistance of Joyce Lawson and Dee Rust of the Department of Radiation Oncology, Stanford University School of Medicine. To all of them we extend our heartfelt thanks. We are particularly grateful to Don Goer of Schonberg Radiation, Inc., and Robert Morton of Siemens Medical Laboratories, Inc., who read the entire manuscript. We have benefited from their many incisive comments and suggestions. We also express our appreciation for the extensive secretarial assistance in prepar-
ing the manuscript and for the drafting assistance in preparing the line drawings. These include: Connie Allen, Dolan Chan, Juanita Clack, Harry Lewis, Hisae Liang, Sumiko Oshima, Sherry Takahashi and Hal Westcott. Finally, we wish to express our deep appreciation to two organizations for their assistance, namely, Varian Associates and Stanford University School of Medicine. They provided facilities for word processing, computer graphics, and reproduction as well as moral support.
Medical Electron Accelerators
C H A P T E R
The Medical Electron Accelerator
OVERVIEW THE NEED FOR MEDICAL ELECTRON ACCELERATORS Cancer is the second leading cause of death in the United States, accounting for one-fourth of all deaths. It is the leading cause of death in the 35-54 age group.' Cancer is not just a disease of the elderly. Approximately one-half of all cancers are detected in people at ages less than 60, and the tumor may have been growing for many years before that. Figure 1-la and b shows cancer death rates for a few selected sites.' Lung is now the most prevalent site associated with cancer deaths in the United States in both males and females. It has been estimated that about 30 percent of all cancers in the United States are caused by smoking. Figure 1-2 shows cancer incidence by site for males and females. About 50 to 60 percent of all cancer patients in the United States will receive radiation therapy,56 as definitive therapy, for palliation, or as an adjunct to surgery. This totals more than 500,000 new radiotherapy patients per year in the United States, plus about 150,000 patients returning for retreatment, for persistent or recurrent disease. In 1986, there were over 1,950 radiotherapy machines in the United States treating an average of 230 new patients per year per machine.20 To serve a world population of 5 billion with the same incidence of cancer and with the same average patient load per machine, about 20 times as many machines would be required as are in use in the United States. About one-half of all radiotherapy patients are treated with curative intent. The remaining one-half are treated for palliation,l6 that is, for relief or prevention of specific symptoms. The median course of treatment for palliation is 10fractions of 300cGy in 15 days. The median treatment for cure is more like 25-35 fractions of 200 cGy in 25 to 35 days for a total of 5000 to 7000 cGy. Of those patients presenting with locoregional disease, 56 percent will be cured.56
work, how optimal designs are achieved, how their components and subsystems are combined to constitute an accelerator, and how the accelerator is integrated into a complete facility and applied in various radiotherapy modalities. Chapter 1 presents a history of electron accelerators and addresses cancer incidence, the goals and rationale of radiotherapy, an elementary description of accelerators, and criteria for optimal accelerator design. This is followed by a survey of isocentric machines that have been introduced by various manufacturers over the past decades and a more extensive comparison of contemporary machine types. Chapter 2 describes various radiotherapy modalities. Chapters 3-11 relate to the theory and design of accelerator hardware. Chapters 3-6 address fundamental aspects of microwave power generation and transmission, microwave structures for acceleration of electron beams, and the associated equipment. Chapter 7 discusses the electron beam inside the accelerator, its generation, acceleration, confinement, bending, and treatment beam formation. Chapters 8-1 1 describe the radiation head, control and safety systems, and the design aspects of multi-x-ray energy machines. Chapters 12-14 relate to equipment and facilities to apply the accelerated beam. These include items in and on the radiation head such as multileaf collimators, the patient support assembly, simulators, treatment planning systems, portal imaging systems, record and verify systems, computer networks, and accelerator usage and maintenance. Chapters 15 and 16 describe microtron and other types of medical electron accelerators. The physics of generation, interaction and application of radiation beams is addressed in Appendix A. This is followed by a survey of historical and contemporary accelerators in Appendix B and by a tabulation of systems of units and a summary of symbols and terminology in Appendix C.
GOALS OF RADIOTHERAPY OUTLINE OF THIS BOOK This book is primarily about megavoltage microwave electron linear accelerators in medicine, about their origins, how they
In addition to trying to save the patient's life, maintaining quality of life is a major goal of radiation therapy. Figure 1-3 compares quality of life outcomes of two cancer patients; one
1
2
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
2
AGE-ADJUSTED CANCER DEATH RATES' FOR SELECTED SITES, MALES, UNITED STATES, 1930-1988
0 1 1930
1 1 1940
1 1 1950
:
1 1 1 1 1 1 1 1 1 1960 1970 1980 1990 Year
AGE-ADJUSTED CANCER DEATH RATES* FOR SELECTED SITES, FEMALES. UNITED STATES. 1930-1988
Year
'Adjusted to the age dtstrlbutlon of the 1970 US Census Populat~on
'Adjusted to the age distribution of the 1970 US Census Population.
Sources of Data US National Center for Health Statlstlcs and US Bureau of the Census.
Sources of Data: US National Center for Health Statistics and US Bureau of the Census.
\
FIGURE 1-1
\
. Cancer death rates (from Ref. 7) (left) Male. (right) Female.
treated by surgery, with gross loss of function; the other treated by radiation therapy.11 Other examples of the use of radiotherapy for preservation of function are to conserve the female breast (for cosmesis); prostate (for better sexual function); bladder, (for more convenient urination). Although quality of life is a major factor, other factors are more widely used to compare efficacy of alternative treatment protocols. Freedom from recurrent disease is used as a basis for comparison of efficacy, but the most widely used measure is survival, the most frequently used being 5-year survival. There have been improvementslo in 5-year survival over the past three decades in treatment of some sites, as shown in Table 1-1. Examples of such sites having relatively high cancer incidence are breast, bladder, and prostate. Improvements in imaging equipment and medical training have led to a more rational selection of patients to treat for cure versus palliation. Although there have been no clinical trials to demonstrate unequivocally that choice of equipment affects cure rates, national surveys
show that the most experienced medical teams get the best results and they tend to have the best medical accelerators.
IMPACT OF TREATMENT COURSE FRACTIONATION ON MACHINE PERFORMANCE REQUIREMENTS The standard course of treatment is divided into daily fractions of about a 200-cGy tumor dose, delivered over a period of, for example, 6 weeks for a 6000-cGy cumulative dose. Fractionation exploits the difference in response between normal and cancerous tissue in their recovery from small doses of irradiation. Normal tissue is better organized and regenerates better than the tumor after a radiation insult. (Think of how quickly tissue regenerates at a cut in your finger.) Fractionation also impacts on requirements for machine performance. Some aspects of fractionation response are as follows56:
3
OVERVIEW
1989 ESTIMATED CANCER INCIDENCE BY SlTE AND SEX' Melanoma of Skin 3% Oral 4% Lung 18% Pancreas 2% Stomach 3% Colon & Rectum 14% Prostate 23% Urinary 10% Leukemia & Lymphomas 9% All Other 14%
\
Melanoma
3% Pancreas 14% Rectum
Leukemia & 7% Lymphomas 11% All Other
.ExcIudong nonmelanoma s k ~ ncancer and carelnoma on sotu
\
\
1989 ESTIMATED CANCER DEATHS BY SlTE AND SEX Melanoma of Skin 1% Oral 2% Lung 34% Pancreas 4% Stomach 3% Colon & Rectum 11% Prostate 12% Urinary 5% Leukemia & Lymphomas 9% All Other 19%
FIGURE 1-2 Ref. 7).
Melanoma
5% Pancreas 12% Rectum
Leukemia & 9% Lymphomas 19% All Other
. Cancer incidence and deaths by site and sex (from
1. Repair of sublethal radiation damage to cells occurs in both normal and cancerous cells. This repair is completed within a few hours after irradiation. Consequently, about 6000 cGy is required in 30 daily fractions to sterilize the same number of cells as 2000 cGy in one exposure. An accidental overdose of say 2000 cGy is far more damaging than just its 10:l ratio to a 200-cGy dose. The implications for dose limiting interlocks for patient safety are obvious. 2. Regeneration of tumor cells can result in some tumors speeding up their growth rate by as much as a factor of 10 after initiation of a course of radiotherapy. This regeneration can require that the course of treatment not be extended excessively or interrupted unintentionally. Unplanned extended machine downtimes can be detrimental to the patient. 3. Redistribution of cells after each irradiation refills the more radiosensitive phases of the cell division cycle, especially with the widely varying rates of progression of tumor cells through the cell cycle. This redistribution occurs faster in some tumor cells than in normal tissue cells and can be exploited by treating all portals in each daily
-
FIGURE 1-3 Surgery versus radiotherapy for cancer of the mouth (from Ref. l l a , Fig. 7.4b and 7.8b).
fraction and, for some tumors, by hyperfractionation, treating the patient more than once per day. These procedures create a need for ergonomic machine design to facilitate rapid positioning of the machine and patient so that a large daily patient load can be maintained. 4. Reoxygenation of previously hypoxic and hence radioresistant tumor cells increases their radiosensitivity by a large factor. Selective loss of radiosensitive cells near the blood supply facilitates diffusion of oxygen from the .blood vessels to these previously hypoxic tumor cells. Because of reoxygenation with conventional fractionated x-ray and electron therapy, improvements in therapeutic ratio with heavily ionizing radiation (e.g., neutrons) have not been as dramatic as was once hoped. Hence, accelerators producing megavoltage x-rays and electrons remain the mainstay of radiotherapy.
USER PREFERENCES FOR BEAM MODE X-ray energies Table 1-2 shows the caseload mix for one large metropolitan area and the percentage of physicians who preferred a particular modality for treatment of each tumor site. The range of tumor sites led physicians24 and physicists46 to recommend use of a low megavoltage x-ray mode for the majority of cases but a widely separated high energy x-ray mode for about one-fourth of the patients and an electron mode for about one-eighth of the patients. As shown in Tables 1-2 and 1-3, the choice of x-ray energy for particular tumor sites is based on a number of factors including depth dose, penumbra, buildup in superficial layers, and buildup at air cavities. Ahigh energy x-ray mode provides a clear advantage in treating large tumors in thick sections of the body such as the lateral pelvis, but also in protecting bone near the skin because of slower
4
TABLE 1-1
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
. Cancer incidence by site-improvements
in 5-year survival
New cases in 1987 Site
Thousands
percent
Five-year survival (whites) 1960-1963(%)
1977-1983(%)
Points
Brain and C N S ~ Oral and Pharynx Breast (fem) Lung Esophagus Stomach Pancreas Ovary Colon Rectum Bladder Prostrate Leukemia Other blood and lymph Other Total Average "Central nervous system (CNS).
buildup of dose versus depth, such as in treating some tumors of the head. Meredith49 discusses the variation of the shielding effect of bone and dosage in soft tissue as functions of x-ray photon energy and the corresponding optimal range of x-ray beam energy.
Electrons There are several sites where the tumerous tissue overlies very radiosensitive normal tissue. An example is the treatment of chest wall tumors while protecting the lung by irradiating the patient with electrons of the appropriate energy and hence limited range. Electrons are also used, for example, to produce a boost dose after x-ray therapy, such as to a tumor lying near one side of the head or neck.
TREATMENT BEAM GENERATION The primary application of electron accelerators in medicine is in the treatment of cancer. A tumor can be detected by diagnostic x rays, which are produced by accelerating electrons to about 0.1 MeV (million electron volts) (see Appendix C for definition and units) and directing them onto a metal target. X rays of this energy are attenuated by a large factor, of order 100, in passing through the body, and attenuation is greater in bone than in flesh. Very small thickness, density, or atomic number differences produce detectable changes in x-ray transmission. Thus, anatomical features are readily seen
in the x-ray diagnostic film or other display. However, to treat the tumor, megavoltage x rays, typically in the range of 4 to 25 MeV, are used because they are attenuated by only a factor of about 2 in passing through the body. In addition, the energy deposition (or absorbed dose) builds up over the first 1-4 cm of penetration, so the sensitive layers of the patient's skin receive only a fraction of the dose at depth. Thus, the megavoltage x-ray beam can be aimed at the tumor from a number of directions, producing a crossover of high radiation dose at the tumor without producing a harmful dose in the rest of the patient's body. These megavoltage x rays are produced by accelerating an electron beam to millions of volts of energy and directing this beam onto a metal target. For tumors that overlie especially radiation-sensitive tissue, such as the spinal cord, the patient is treated with the megavoltage electron beam directly, without converting to x-rays in a metal target. Electrons penetrate the patient's body only to a distance in centimeters of about one-half of their energy in megaelectron volts (e.g., 5 cm for a 10-MeV electron beam). In diagnostic x-ray tubes, the electrons are accelerated to about 0.1 MeV across a single gap. A high-voltage transformer at a power line frequency of 50 or 60 Hz (Hertz, cycles per second), is used to develop the accelerating voltage (see Fig. 1-4). Since many millions of volts cannot be held off readily by a single gap, electrons in microwave linear accelerator type radiotherapy machines are accelerated to megavoltage energy by passing them through a succession of gaps (see Fig. 3-3b). Each gap is at the center of a microwave
5
OVERVIEW
TABLE 1-2
. Caseload and beam preferences-Philadelphia
area Physician's beam preferences Low X
Caseload(%)
Body area
Lung Pelvis Prostate Cervix Head and neck Breast (intact) Adomen Pancreas Brain primary Chest wall Trancheaandesophagus Nodes Bone mets Brain CNS mets Other
22 20
High X
Electrons
35 17 4 83 96
7 7 5
0 74 52 78
4 3 2
3 18 3 6 100
Beam utilization at two multimodality departments Low X(%)
High X(%)
Electrons boost(%)
71
23
12
resonant cavity (see Fig. 1-5), which is powered at a frequency of about 3 billion hertz. Because the cavity resonates at this frequency, the required drive power from the magnetron or klystron radio frequency (rf) source is only about 0.01% of the oscillating reactive power in the cavity. In simplified analogy to the diagnostic x-ray system, each cavity acts like a small transformer. Oscillating rf current flows on its internal surface and acts like a single turn primary winding; the electron beam bunch flowing synchronously across the gap acts like a single turn secondary winding. Each gap in turn TABLE 1-3
Electrons alone(%) 6
adds to the energy of the electron beam bunch. For example, a beam energy of 20 MeV can be obtained in a series of 28 cavities, gaining about 0.7 MeV per cavity (gap). Yet the entire multicavity structure is at ground potential. The only high voltages are at the electron gun (15-35 kV), which initiates and injects the electron beam and at the magnetron or klystron (50-150 kV), which provides the microwave power to excite the multicavity accelerator structure (see Fig. 1-6). A few megawatts (MW) of microwave peak power are required to excite the accelerator structure to suitably intense
Importance of beam characteristics versus site Beam characteristics
Site
Brain Head and neck Breast Thorax Lymphoma Pancreas Whole pelvis Pelvic cone Down Pediatrics
Depth dose
1 1 1 1 1
Optimum energy vs. site
Build-up
Beam penumbra
Bone dose
1 1 1 1
1 1 1 1 1
1 1 1
1
1
1
Neutron
60Co
. .
4 MV
6 MV
b
..
10-15MV
1
•
+
. 4
4
1
.
.
1
1 1
+-
~ 1 8 W
b
b
. b
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
6
12 Rectifiers (6 pulseslcycle)
12 Rectifiers (6 pulseslcycle)
6 Rectifiers (6 pulses/cycle)
A
A
X-ray Tube Voltage +
It--1 cycle ---)I
I cycle
+
Transformer Windings: Seconday
FIGURE 1-4
. Diagnostic x-ray tube high-voltage circuits.
accelerating fields but only a kilowatt (kW) or less of average power is required to produce adequate intensity in the x-ray treatment beam. Thus, a pulse modulator is used to pulse the magnetron or klystron on for a few microseconds (ps) every few milliseconds (ms) several hundred times per second.
HISTORY OF ELECTRON ACCELERATORS DIRECT ACCELERATORS Laughlin47 presents a concise history of the technological development of radiation therapy. The earliest machines were direct accelerators, in which the entire voltage was held off over a long insulating column. Cockroft Walton generators, M a n generators, resonant transformers12 (Resotrons), Van de Graaffs, and Dynamitrons are examples of this technique. All of these machines were large and cumbersome and difficult to move around a patient. Their energy was also limited to about 2 MeV in orientable machines.
f
i
Current -I
FIGURE 1-5
- Microwave accelerator cavity circuit.
Electric Field
BETATRONS The betatron44357+".68 was invented in 1940. It employs a transformer technique of magnetic induction. The electrons travel around in a circle, many thousands of times, acting like the electrical current in the secondary winding of a transformer, hence multiplying the alternating current primary voltage of about 10 kV (kilovolts) to perhaps 2 5 4 5 MeV. The electron beam is confined within an evacuated donut by the weak focusing forces of the betatron magnetic field as it rises in intensity in synchronism with the rise in electron energy from injection to full energy in one-fourth cycle of a sine wave. The beam current
7
HISTORY OF ELECTRON ACCELERATORS
Accelerator Structure Modulator and High Voltage P.S.
>
A
f
E'FLr Accelerator A
RF Subsystem
Primary Power Distribution
4
Target
Accessories
Water Cooling Subsystem
A
I -
t
Collimator Drives & Position Sensing
:
Gantry Drive 8 Position Sensing
Y
A
Subsystem
Wlfe:~cJe
Dosimetry
4
To Accelerator Structure
Y -+
L F
'..
A
h
-
CO'limatOr --+ -+Assembly
Relay Interlocks (In Clinacs)
-_ v
Console and Controller
-
ETR (Couch) Drive & Position Sensing
A 7
4
-
FIGURE 1-6
. Block diagram for low-energy machine.
confined by this technique is relatively small, so the x-ray intensity and flattened field size of betatrons is quite limited. The electron beam could also be extracted from the circular orbit at any desired energy by pulsing a deflection coil at the appropriate time within the one-fourth cycle sine wave rise of the main magnetic field. Betatrons of 25 to 45 MeV are quite large and heavy, so they were never developed for gantry type isocentric mounting around a stationary patient. Instead, the patient table horizontal axis was moved in a circle as the betatron was rotated, lifting the patient quite high to permit treating from below. Smaller low energy betatrons, which could have been isocentrically mounted, had an x-ray intensity that was too low at the desired distances. The electron beam current in betatrons was just too low for modem clinical x ray use.
MICROWAVE ELECTRON LINEAR ACCELERATORS (LINACS) The term "linac" is a contraction of the term linear accelerator. It means that the charged particles travel in a straight line as they gain energy from the accelerating electric field. The term "linear" is used to distinguish from other types of particle accelerators, such as the cyclotron, in which the particles travel in a spiral, or the betatron, in which the particles travel in a circle. Because the linac has superseded direct accelerators and betatrons and is displacing 60C0, the history of inventions and developments leading to the modem linac are presented in the following subsections in some detail. In the decade from 1935 to 1945, the basic theoretical understanding and practical ap-
plication of microwave devices and systems was developed, including invention of the microwave cavity and of the klystron and magnetron tubes as sources of microwave power. From 1945 to 1958, linacs were invented and built by physics research laboratories, and the multimegawatt klystron was developed. British industry built magnetron driven 4-MeV isocentric medical linacs of limited rotation and also a few nonisocentric medical linacs of higher energy. In this same period, in the United States, Stanford University built a 6-MeV orientable medical linac and a few stationary medical linacs of higher energy for electron therapy,48,68,73,* which were all driven by klystrons. In 1962 in the United States, Varian Associates introduced a fully rotational isocentric 6-MeV bent beam linac. The side coupled standing wave accelerator structure was invented at Los Alamos National Laboratory in the late 1960s and Varian Associates applied it in introducinga fully rotational isocentric 4-MeV in-line linac in 1970. Several other manufacturers initiated manufacture of medical linacs during this period. Since 1970, machine performance characteristics have been extended and precision has been improved. Some examples are the improved accelerator structures and bend magnet systems and the introduction of microwave techniques to achieve widely spaced dual x-ray energies. This accelerator history is expanded upon in the following subsections.
Microwave cavities In the mid-1930s Hansen32.33, then at the Stanford University Physics Department, wanted to accelerate electrons to millions
8
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
of volts by magnetically reflecting them back and forth through a hollow copper cavity filled with high-power ultra-high-frequency radio waves called microwaves. He developed the mathematical techniques for analyzing such cavities in order to find optimum geometries and found that the efficiency could be a hundred times greater than with the resonant circuits of that time, which might use capacitors and coils. The physics department agreed on the name Rhumbatron for Hansen's resonant cavity invention, from the Greek words Rhumba meaning rhythmic oscillations and tron for the place where they occur. Depending on its application, the rhumbatron might look like a tomato can or like a soft tennis ball with opposite sides pushed in. Hansen tried to build an electron accelerator employing a rhumbatron and powering it with the only device of the time that could provide power at such high frequencies, a little "acorn" tube. It had to be little to work at such high frequencies and hence it could not produce much power. Hansen needed a totally new method of providing microwaves to obtain enough power to accelerate electrons to millions of volts. Fortunately, within the following 5 years, two new types of powerful microwave sources were invented: the klystron in 1937 and the high-power pulsed magnetron in 1939.
Klystron invention While a graduate student at Stanford in 1934, Russell Varian developed a close association with Hansen. Russell's brother Sigurd Varian was a pilot for Pan American Airways. In 1936 cities were being bombed by air in Spain and China and Sigurd recognized the need for a way to detect enemy aircraft through cloud cover. It occurred to Russell that Hansen's rhumbatron might be adapted in some way to generate the microwave power needed to detect aircraft. The two brothers had started their own private research laboratory but soon recognized that it was too limited for the task, so in 1937 they negotiated an arrangement to carry on their work at Stanford University. They were appointed Research Associates without salary and were given a budget of $100 for materials and supplies. Russell conceived a number of ideas for generation of microwave power using Hansen's rhumbatron. In developing a classification for all the schemes he thought of an idea that did not fit any of the classifications-the velocity modulation principle. A small oscillating signal in one rhumbatron varies the speed of a steady stream of electrons. As these flow downstream the speeded up electrons catch up with the slowed down electrons, forming bunches of electrons that pass through a second rhumbatron in which the kinetic energy of the electron stream could produce high power. The idea25.74.75 was conceived on July 22, 1937, and Russell's brother Sigurd built a tube and got it working within 4 weeks. The beauty of this invention was that individual separate large components (cathode, rf input section, rf output section, beam collector) could be used and still produce high power at very short wavelengths, 10,000 times shorter than the ordinary radio
waves of local radio stations. This was a major breakthrough, a complete escape from the previous technical constraints requiring that microwave sources be physically small, and hence, be of low power. The highest pulsed output power achieved with a klystron during World War I1 was 30 kWvery high indeed for that time but still too low for use in accelerators.
Magnetron invention The British were also concentrating heavily on a different form of microwave power generating device, called a magnetron. This is a circular beam microwave oscillator, in contrast to the klystron, which is a linear beam microwave device that can be operated either as an amplifier or an oscillator. In the magnetron, a number of cavity resonators are arranged in a circle and a magnet bends the electron beam so that it forms a number of curved spokes that sweep around to excite power in each resonator. This multicavity concept was suggested by research workers in the United States in 1936, England and Germany in 1938, and Russia in 1940, but the maximum power achieved was only about 100 W. The wartime need of radar for a transmitting tube capable of producing very high pulsed power at 10 cm or less (a 10 cm wavelength was larger than absorbers in the atmosphere but small enough to be transmitted in a narrow beam from an antenna of practical size) led the British laboratory workers and the British industry to develop a tube with about 10,000 times the pulse power output that had previously been available. It provided spatial resolution suitable for distinguishing not just whether an airplane was detected but whether it was friend or foe. J. T. Randall and H. A. H. Boot,6 University of Birmingham; England, invented the first such tube and in 1939 they achieved over 0.1 MW at 10-cm wavelength. Further development increased the pulsed output power to over 1 MW. Finally, here was a tube suitable for powering a modest energy microwave linear accelerator, but that application had to wait until after the war ended.
Microwave linac invention With the availability of high-power magnetrons and microwave techniques and components after the radar developments of World War 11, about 10 groups independently started inventing and building microwave electron linacs. Two of these groups became preeminent: W. W. Hansen's group at Stanford University and D. D. Fry's group at the TelecommunicationsResearch Establishment (TRE), Great Malvern, England, which subsequently became part of the Atomic Energy Research Establishment (AERE), Harwell, England. These two groups leapfrogged each other's accomplishments for the first few years but had limited knowledge of eachother's work until mid1947. Fry's group22,23,53.61.77completed the theoretical design of a 45-cm long, 0.5-MeV section of accelerator guide in September 1946 and accelerated electrons for the first time in late
HISTORY OF ELECTRON ACCELERATORS
November 1946. Totally independently, Hansen'sgroup26completed their work on a 90-cm accelerator guide powered by a magnetron and obtained 1.7 MeV in early 1947. Fry's gro~p22~sO proceeded with the development of a 2-m long accelerator guide driven by a magnetron and achieved 3.5-MeV of electron energy in ~ o v e m b e 1947. r Hansen's group4-14extended their accelerator guide to 3 m and obtained 4.5 MeV by November 1947, and subsequently 6 MeV, still using a magnetron. Fry22 states that Cutler18 in 1944 was the first to solve the mathematical equations, which showed how a corrugated cylinder could be designed to match the phase velocity of the traveling electromagnetic wave to the electron velocity for application to traveling wave tubes. The corrugations are copper disks and Woodyard79 in Hansen's group suggested using such a waveguide as a linac forelectrons.
Multimegawatt klystron invention E. L. Ginzton started working with W. W. Hansen and the Varian brothers in 1938 and conceived a set of ideas in 1944 that led to a proposal in 1947 to build a klystron 1000 times more powerful than the most powerful wartime klystron. In cooperation with Chodorow et a1.,15 a successfi~lhigh-power klystron was first demonstrated in 1949, and after three more years of effort, the original goal of 30 MW, 1000 times the highest wartime klystron power, was achieved. This development opened the way to building an electron linear accelerator of 1000 MeV for physics research, and eventually, very compact medical accelerators at energies such as 25 MeV.
First stationary linac for radiotherapy With the achievement of 3.5 MeV with a 2-m accelerator section by Fry's group, collaboration of three groups was arranged under the auspices of the British Ministry of Health toward the end of 1948 to build an x-ray linac for clinical use. These groups were the Radiotherapeutic Research Unit of the Medical Research Council (MRC) under Dr. L. H. Gray, the linear accelerator team of AERE under D. W. Fry, and the Metropolitan Vickers Electrical Company (later renamed Associated Electrical Industries), with C. W. Millerso as project leader. Installation commenced at Hammersmith Hospital, London, in June 1952 and the first patient was treated on August 19, 1953. The machine employed a 2-MW magnetron and a 3 m stationary accelerator guide with a rotatable 90" magnet on the end. It was typically operated at 100 cGy/min (100 radlmin) with the 8 MeV x-ray beam flattened over maximum field sizes of 25 cm in diameter or 15 X 20 cm rectangular. The treatment room floor could be moved vertically and the treatment table moved laterally as the 90" magnet radiation head rotated in order to provide patient portals over a range of angles. These motions were coordinated so as to rotate the x-ray source on a path of constant distance from a
9
selected point in the patient. This was called pendulum therapy and was used as well with betatrons.
First orientable linacs for radiotherapy While the 8-MeV linac for Hammersmith was being built, the Ministry of Health agreed to have a number of 4-MeV linacs built for radiotherapy centers in England. Machine designs and specifications were developed by the Medical Research Council (MRC) and AERE, and were supplied to British industry. The design of an isocentric gantry mount for the accelerator guide was conceived by P. Howard-Flanders36.37 at MRC, Hammersmith in 1949. The x-ray beam was emitted along the axis of the accelerator guide, which was 1 m long and was designed by AERE. The first double gantry unit19 was installed at Newcastle General Hospital in August 1953. The first single gantry units] was installed at Christie Hospital, Manchester, and operation was started in October 1954. The single gantry machine could be rotated over an arc of 120°, from 15" beyond vertical to 15" beyond horizontal, or to 30" beyond horizontal by lowering part of the treatment room floor (Fig. 1-7). Field sizes to 20 X 20 cm at 1 m were provided at a typical dose rate of I00 cGy/min. Independently, a program of radiotherapy linac development proceeded in the United States. Conceptual work was initiated in 1950 by Kaplan and Ginzton. The first machine27 was built in the Stanford University Microwave Laboratory, with three faculty members and seven graduate students participating. Installation78 of the machine in Kaplan's radiology department in San Francisco was started in 1954. The accelerator guide (Fig. 1-8) was built as a sealed-off vacuum tube, without any vacuum pump. This avoided the problems of gun and accelerator guide contamination from the oil diffusion pumps of that time and avoided the complexity of a rotating vacuum seal between the orientable accelerator guide and the oil diffusion pump, which had to be kept vertical. The 1.65-m long accelerator guide was electroformed by depositing copper
FIGURE 1-7 . Firs, orientable linearaccelerator-Theorthotron (from ~ , f 51). .
10
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
Accelerator
FIGURE 1-8
X- Ray Target
. First Stanford 6-MeV clinical Accelerator guide (from Ref. 27).
in an acid bath onto an aluminum spacer copper disk mandrel, then etching out the aluminum. During the construction pmgram the high-power klystron14 was developed at Stanford, so a 1-MW version was used to power the accelerator. Some accelerator guides were built with a gold target sealed into the end. Others were built with a thin titanium window, permitting radiotherapy directly with electrons or with x-rays from an external gold target. Field sizes to 15 X 15 cm were provided at an unflattened dose rate to 110 cGy/min at 1 m. For tumor localization and treatment simulation, a 100-kVp (kilovolt peak) rod anode x-ray tube could be inserted near the linac target position to permit viewing the patient portal with an image intensifier as well as for taking portal films. The accelerator with an in-line radiation head was installed in a trunnion mount normally used with Van de Graaff accelerators, permitting vertical travel of 150 cm and slightly more than 90" accelerator orientation from vertical to approximately horizon-tal around a reclining, standing, or seated patient. The very precise, sharp, intense, deeply penetrating x-ray beam from this linac gave Kaplan et al.41 the confidence to accept an infant with retinoblastoma as the first patient for treatment, in January 1956. The retina was irradiated without damaging the lens or cornea of the eye and this patient was still doing well 32 years later (in 1988), with his vision in the treated eye intact. The electron beam was also used directly to treat patients with superficial lesions. In the period around 1958, Varian Associates was designing and building linacs for a number of applications, such as physics and chemistry research, food irradiation, sterilization of medical supplies, and radiography of thick objects. It also built at this time a rotatable system of magnets69 which transported and scanned an electron beam from a 50 MeV linac (accelerator guide built by Stanford), providing electron therapy of a reclining patient from any angle around a 360" circle.
A team had been assembled with accelerator related experience from high-energy physics laboratories and other sources. This was later supplemented by engineers having experience with British linacs. In response to the advocacy of E. L. Ginzton, a program was initiated in 1958 to develop and manufacture a clinical linac. With the British industrial work and the Stanford University work as a starting point, this team, under the direction of C. S. Nunanz8 developed the designs for a 6-MeV isocentric linac so compact that full 360" rotation about a reclining patient could be achieved. This would permit anterior and posterior irradiation of a supine patient. A novel isocentric patient table was developed, which provided clearance for the radiation head and permitted fully lowering the patient to the floor. The first production machine (Fig 1-9) was installed in 1962 at the Stanford University School of Medicine in its new facility in Palo Alto. This machine2-3.30,55employed a 1.5-m
FIGURE 1-9 . ~ j n 360° t isocentric linear accelerator-The Clinac 6 (multiple exposures) (from Varian).
11
HISTORY OF ELECT'RONACCELERATORS
Water Cooled Fixed Target
Electron Window
-
I
Flattening Filter And Dual Ionization Chamber Assembly (Retracts For Electron Extraction Operation)
FIGURE 1-10 Beam bending system with separate paths for electron and x-ray modes (from Ref. 3).
long accelerator guide mounted horizontally in a gantry and driven by a 2-MW magnetron. A unique 90" magnet system (Fig. 1-10) provided beam energy discriminationby intercepting a portion of the dispersed beam on the edge of a gold target for x-ray therapy. The magnet system redirected the accelerated beam through an electron window for electron therapy without moving the x-ray target. The stack of machined copper parts forming the accelerator guide was silver brazed in a vacuum furnace, thereby also degassing the metal. Small drilled holes in each cavity provided access for tuning and provided highvacuum pumping speed. The accelerator guide was enclosed in a stainless steel cylinder and sealed off. Its vacuum was maintained by a small device called a VacIon pump31, which had been invented in 1956 at Varian Associates for an entirely different purpose. This sputter ion pump provided a clean oil-free high vacuum and it could work in any orientation, so it was ideal for the clinical accelerator. It provided a major step forward in the reliability of these machines.
apertures were much larger than the beam diameter, which with other factors limited the shunt impedance (47 and 56 M Wm (megaohm per meter) in the machines of Figs. 1-8 and 1-9, respectively). In 1968, Knapp et a1.45 invented the side-coupled standing wave structure, in which the microwave power is coupled between axial cavities via slots to side cavities, and a small aperture is provided on the axis for passage of the beam (see insert of Fig. 1-15). This separation of functions permitted designing the slots to the side cavities for desired coupling and designing the axial cavity geometry in a rather spherical shape with reentrantnoses and small beam aperture, attaining a shunt impedance at 10 cm operating wavelength of about 80 M Wm in initial designs and as much as 110 M Wm in modem 2 (90" phase shift per axial accelerator guides. In the ~ 1 mode cavity and 90" per side cavity) the side cavities are at nulls 180" apart along the standing wave so they dissipate negligible microwave power and the axial cavities are at the maxima of the standing wave. Neighboring resonant modes are spread far apart, an important frequency stability criterion. (The nearest neighbor resonant modes are about lMHz (megahertz) away in resonance from the fundamental mode and they have a cosine-like distribution of electric field intensity over the guide length-high at one end, zero in the middle, and reversed high at the other end of the length.)
In-line standing-wave linacs for radiotherapy The invention of the side coupled standing-wave accelerator structure permitted elimination of the bend magnet and use of an extremely short in-line accelerator guide in a 360" isocentric gantry for low megavoltage radiotherapy machines. Varian Associates applied this technique, along with miniaturization of the electron gun and target, developing the first-in-line 360" isocentric machine43.62 Figure 1-11, a 4MeV unit with an 80-cm source to axis distance, in 1969
Standing-wave accelerator guide For a given beam energy and microwave power loss in the cavity copper surfaces the required length of the accelerator guide is inversely proportional to shunt impedance (a term that defines the efficiency of conversion of microwave power to gain in beam energy per unit length) (gain). The early medical linacs employed a corrugated (disk loaded) accelerator guide in a traveling-wave mode (Fig. 1-8). The aperture in the disks served two purposes, to couple microwave power from cavity to cavity along the accelerator guide and to permit passage of the beam along the axis. To provide proper coupling, the
Primary Collimator
O~ticsand Source lsocenter
.
Range Finder Optics and Source
FIGURE 1-11 First 360' isocentric in-line linear accelerator-The Clinic 4 (from Varian).
12
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
VARIAN
KLYSTRON
FIGURE 1-12 . Multimode linac with microwave energy switch and 270" doubly achromatic magnet-the Clinac 1800 (from Varian).
and the first-in-line 6 MeV, 100 cm, 360" isocentric machine in 1977.
Bent beam standing-wave linacs for radiotherapy Although distributed doubly achromatic (energy-independent focal properties) magnet systems were known in the 1950s, the theory of doubly achromatic magnets compact enough to use in clinical linacs had not yet been developed. Machines with a nonachromatic magnet system (such as the 90" system in the above isocentric 6-MeV machine produced in 1962) or a singly achromatic magnet system (axial rays of differing energies from the accelerator converge over a spread in angles at the x-ray target) have one basic problem. Small changes in mean energy of the accelerated beam energy spectrum result in changes in mean angle and position of the beam with respect to the axis of the conical flattening filter after bending in the magnet. This causes variations in treatment field symmetry, which are difficult to compensate completely by feedback systems from symmetry monitors. This latter problem becomes especially severe with the narrow primary x-ray lobe of higher energy machines. One solution is to use a doubly achromatic bend magnet, usually a 270" magnet to minimize isocenter height. The distribution of positions and angles of all the rays making up the electron beam from the
accelerator are reproduced at the output of a doubly achromatic bend magnet independent of the energies of these individual rays. For machines of energy higher than 6 MeV, this solution is used; that is a doubly achromatic 270" magnet (Fig. 1-12) (see Chap. 7 on magnets for further discussion) is installed in the radiation head. Figure 1-13 shows a compact doubly achromatic 270" magnet proposed by H. A. Engez* in which the magnetic field gradually increases and then decreases along the beam path.42 The different energy rays in the electron beam at the input of the magnet all coalesce at the output. This magnet 270'
MAGNET
MAGNET POLES
SECTION d l - d 2
FIGURE 1-13 Ref. 42).
. Beam paths in 270" doubly achromatic magnet (from
13
HISTORY OF ELECTRON ACCELERATORS
-AL2 (mil)
- 80
- 60
+ A L o (mil)
FIGURE 1-15 . Microwave energy switch for control of Axial field ratio EzIEo (from Ref. 70.).
FIGURE 1-14 . Machine with electron applicator-The Clinac 20 (from Varian).
is difficult to manufacture and align so more practical 270" magnets were developed such as a type in which the magnetic field steps from low to high and back to low along the beam path.10 Figure 1-14 shows a 20-MeV machine in the treatment room, set up for electron therapy.
Dual x-ray energy standing wave linacs Compact techniques (e.g., see Figs. 1-12 and 1-15) have now been developed, which provide both high and low megavoltage x-ray treatment capability in the same machine. In these units, the field level is maintained for proper electron bunching and acceleration in the early part of a standing wave guide but is raised or lowered in the later part of the guide. One way to accomplish this is by moving the posts in one side cavity in order to change the ratio of coupling fields at its two ends while maintaining its resonance in the 7~12mode, thereby changing the ratio E21Eo of electric fields in the axial cavities. This permits switching between two widely displaced x-ray energies?' such as 6 MeV for head and neck and 18 MeV for pelvic tumors. Switching to the lower energy is also an advantage
when taking port films or electronic portal images. In addition to increasing beam performance capabilities, a number of treatment aids and accessories have been developed,29 including patient immobilization and positioning systems, electron arc therapy, and treatment recording and verification systems.
The pioneers From its earliest application by clinicians, which was really a spin-off from programs in basic physics research, the microwave electron linear accelerator in the ensuing decades has become the machine of choice for cancer therapy. The designs of these modern machines did not just appear out of thin air. Historically,they are rooted in the creative and persistent efforts of a few research physicists, engineers, and radiotherapists.
RECIRCULATING ELECTRON ACCELERATORS By using magnets to recirculate the electron beam through the microwave accelerator cavity (or cavities) one or more times, a high-beam energy can be achieved with a low energy accelerating section. After each orbit in the magnet the electron bunch must arrive in phase with the accelerator microwave field. The magnet system acts as an energy spectrometer, limiting the electron energy acceptance to a narrow energy width and consequently limiting to some extent the beam
14
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
Electron Gun
n
d S.W. Linac
I
L - - - - - - - - - - - J
Reflection Magnets
FIGURE 1-18
FIGURE 1-16
. Circular microtron orbits (from Ref. 78a).
current. Three general forms of recirculation have been applied in machines for radiotherapy; the circular orbit microtron (Fig. 1- 16), the racetrack microtron (Fig. 1-17), and the Reflexitron (Fig. 1-18). The history of these devices is reviewed in the following subsections.
Circular orbit microtron In a classic paper in 1944, V. I. Veksler76 at the Lebedev Physical Institute in the USSR proposed the circular orbit microtron. It employs a single accelerating cavity in the gap of
. Reflexitron (from Ref. 58).
a large circular magnet. Each time the bunch of electrons passes through the accelerating cavity it gains an increment of energy and follows a correspondingly larger diameter circular orbit. All the orbits are tangent to each other at the accelerator cavity. Typically, the length of each orbit is one wavelength longer than the previous orbits so the electron bunch arrives back at the accelerating cavity in phase with the accelerating field. The electron beam can be extracted at each incremental energy by moving a magnetic shunt into the path of the corresponding circular orbit. The Canadians34.-58+59were the first to demonstrate this principle in 1950. The USSR laboratories worked intensively on the idea starting in the late 1950s, primarily in the Laboratory for Physical Problems of the Nobel Laureate, Pyotr Kapitza.40 A Swedish company, Scanditronics, was the first to build practical forms of conventional microtrons for cancer therapy, starting in the early 1970s. Because of the very narrow energy spread of the electron beam, it is feasible to transport the beam from a centrally located microtron to two or more treatment rooms by relatively small focusing and bending magnets. The gantry in each treatment room is quite compact, containing only the beam transport magnets and the radiation head. A few dual treatment room circular orbit microtron facilities have been built.
Racetrack nlicrotron
' - 2cb 1 - L 180' "Extract ion" ,/ ~agnet ' "Chicane" Injection
180" "Injection" Magnet
0
CM
E l e c t r o n Gun
FIGURE 1-17
. Racetrack nlicrotron orbits (from Ref. 60).
50
The single accelerating cavity of the circular orbit microtron can be replaced by a linac structure of several cavities by arranging a space for it that is free of magnetic field so that the beam does not curve through the acceleration region. This concept was first suggested by Schwinger in 1946,67 creating racetrack shaped oval orbits with 180" bends in each of two separated magnets. This permits more energy gain per lap, hence fewer orbits for a given energy, smaller magnets, and a more compact machine. Research workers at the University of Western Ontario, Canada developed one form of racetrack microtron.9 Workers60 under the direction of 0. Wernholm at the Royal Institute of Technology, Stockholm, Sweden developed a different form, which produced 50 MeV with a 3-MeV energy gain per orbit.
ELEMENTARY DESCRIPTION OF MEDICAL LINACS
A5O-MeV medical version of this machine has been developed by the Swedish firm Scanditronics.Because of the high energy, the electron beam is scanned in both x-ray and electron modes. The principal virtues of the race track microtron are its compactness for high energy, transportability of the narrow energy spread beam by magnets, ease of changing energy .over a wide range for both electron and x-ray beams, and need for only a relatively low-power microwave source to obtain rather highelectron energies.
Reflexotron The research workers at the University of Western Ontario, Canada, also developed the concept of the Reflexotron.64 Instead of guiding the electrons around a circular or oval orbit around the accelerator structures, they are bent around a much smaller loop and are sent back along the axis of the accelerator structure. This takes advantage of the fact that standing-wave acceleratorscan accelerate in either direction. At a given instant in time the forward flowing bunch of electrons is in phase with the forward force in one-half of the axial cavities and the reflected bunch of electrons is in phase with the backward force in the intervening other one-half of the axial cavities of the accelerator structure. A 25-MeV medical version of this machine has been developed by a Canadian firm, Atomic Energy Corporation, Limited (AECL).72
ELEMENTARY DESCRIPTION OF MEDICAL LINACS In medical linacs, the charged particle is an electron and the rf accelerating electric field oscillates at about 3 billion cycles per second (3000 MHz). For comparison, note that radio waves in the standard AM (amplitude modulation) broadcast channels oscillate at about 1 million cycles per second (1 MHz). The electrons are boiled out (thermionicemission) of a hot cathode (a concave shaped piece of metal heated to over 1000°C) and speeded up in the gun to about one-fourth the velocity of light by a pulsed dc electric field. They are formed (coalesced) into a pencil beam by a convergent electric field between the gun electrodes (See Fig. 4-4). The rf electric field in the accelerating structure then forms the electron stream into bunches, and speeds them up to more than 99 percent the velocity of light, increasing their mass by many times (e.g., by a factor of 13 at 6 MeV). Thus, the electrons become ponderous (massive) and penetrating. It takes about 2.5 cm of water or 0.15 cm of tungsten to stop a 6-MeV electron. Whenever an electron is shaken (i.e., decelerated, deflected, vibrated), it emits radiation. The more violent the shaking, and the higher the electron energy, the harder (more penetrating)the radiation that is emitted. For example, in a light
15
bulb hot filament the electrons are vibrating slowly, have very little energy, and emit quite low energy radiation, some of which is in the visible range. When a 6-MeV electron enters a tungsten target, the individual positively charged nuclei of the tungsten atoms pull on the negatively charged electron, shaking it violently as it passes by one such tungsten atom after another. These high-energy electrons thus emit hard penetrating x-rays, in a forward lobe, giving up a portion of their energy at each successive target atom that they penetrate. As x-rays penetrate the patient they shake loose electrons from the atoms of the patient's tissue, converting them briefly to ions (hence, the term ionizing radiation). These ions can produce further ions in the tissue as they are slowed down (decelerated). The population of ions recombines to form chemically active species. For example, in the presence of free oxygen (0.3 diffused to the cells from the vascular system, water (H20)in the cell nucleus can convert to hydrogen peroxide (H202) and free radicals that can attack the cell deoxyribonuclic acid (DNA) chemically and sterilize the cell so that it will not reproduce. These secondary electrons produced in the patient tissue by x-rays travel primarily forward and produce additional ionization, depositing energy termed a radiation dose. For high-energy x-rays the intensity of this secondary electron flux increases for the first 1 4 cm of tissue, sparing the skin and producing maximum dose at a tissue depth corresponding roughly to the range of these secondary electrons (e.g., 1.5 cm for 6-MV x-rays). Beyond this depth, the x-ray intensity decreases because of absorption and spreading of the beam. By aiming the x-ray beam at the patient's tumor from more than one direction, a cross-over of x-ray beams and summation of irradiation dose can be provided at the tumor while sparing healthy tissue. One way to do this is to position the patient table so that the tumor is at a point in space called the isocenter. (see Figs. 1-19a-c and 1-20). This point is often indicated by wall-mounted and ceiling-mounted laser beams. The radiation beam is then rotated to selected portal angles (or in arcs) around the patient. This technique calls for the source of radiation to be mounted in an isocentric gantry, which is rotated on bearings by a motor drive. The machine radiation head contains heavy metal (e.g., lead and tungsten) radiation shielding to protect the patient from radiation outside the intended treatment beam and thick shielding (concrete and iron) is used in the treatment room walls and ceiling to protect the persons outside. For those patients suitable for curative radiotherapy, the goal is to eradicate the local and regional tumor cells while preserving the function of involved organs and maintaining appropriate cosmesis. All modern microwave electron linacs employ an isocentric gantry. An accelerator waveguide structure is mounted in the gantry, either relatively horizontal, if a beam bending magnet is employed, or normal to the gantry axis, if a beam bending magnet is not required. The term "waveguide" derives historically from the use of a hollow pipe to transport (guide)
16
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
-
FIGURE 1-19 Isocentrically mounted medical linac. (a) With electron applicator. (b)X-ray patient set-up. (c) In-line radiation head.
microwave (e.g., radar) power. Corrugations are used in accelerator waveguides to slow up the waves (somewhat analogous to small jetties at a beach to break up the ocean waves). As a result, the crests of the microwave electric field are made approximately synchronous with the flowing bunches of electrons. Figure 1-21 shows a simplified block diagram of the major parts of a medical linac. The microwave power to accelerate the electrons is provided from a vacuum tube (magnetron or klystron), which is pulsed by a high-voltage modulator. Auxiliary systems provide a high vacuum inside the accelerator guide structure, and cooling and temperature control of its internal conducting surfaces. Insulating gas pressure is maintained inside the rectangular waveguide components, which feed the accelerator waveguide from the microwave power source through ceramic windows. A system of monitors and automatic feedback systems and interlock circuitry main-
tains stable safe operation at values selected at the control console.
MICROWAVE ACCELERATION PRINCIPLE The electromagnetic (EM) field within the open volume of a microwave cavity induces electrical current flow on the inner surface (walls) of that cavity. As the field oscillates, the wall current direction oscillates. During a one-half cycle of EM field oscillation, the electrical current in the walls will flow so as to charge the input end of the cavity with electrons. These electrons on this input end will push the bunch of beam current electrons forward, accelerating the beam, giving it energy. During the next one-half cycle of EM field oscillation, the electrical current in the walls will flow to charge the output end of the cavity with electrons. If the electron bunch were
ELEMENTARY DESCRIPTION OF MEDICAL LINACS 3
UNFLATTENED X-RAY LOBE (18 M'f)
FIGURE 1-20
. Radiation head and x-ray field symmetry (from Ref. 54).
I I
FIGURE 1-21
STAND
GANTRY
. Simplified block diagram of major parts of a medical linac
17
18
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
still in the cavity, it would be decelerated. However, by this time the electron bunch is in the next microwave cavity and is accelerated again because the EM fields oscillation in this next cavity is delayed in phase (time) relative to the first cavity. By stringing a number of microwave cavities together and filling them with EM fields in proper phase relationship to each other, the electron beam bunch can be accelerated in each successive cavity. The subject of microwave acceleration is addressed more completely in Chap. 3.
BEAM CURRENT REQUIREMENTS IN X-RAY MODE The unflattened x-ray intensity on the beam axis from an optimum thickness x-ray target is proportional to the average beam current and to about the 2.6 power of the electron beam energy at the target. This is due to the x-ray lobe cross section (solid angle) being inversely proportional to the square of the electron beam energy and due to the efficiency of conversion of the electron beam power to x-ray power being proportional to about the 0.6 power of the electron beam energy (in the 4 to 25-MeV energy range) (see Appendix A for further details). The required thickness of flattening filter increases with an increase in the diameter over which the field is flattened and increases with x-ray energy. For example, for a field flattened to a 50-cm diameter at 100-cm source axis distance (SAD), which is typical for machines with 40 x 40-cm rated field size (hence, with clipped or unflattened corners), the transmission of the flattening filter on the axis varies by about the - 0.8 power of x-ray energy. Thus, the flattened x-ray intensity is proportional to about the 1.8 power of x-ray energy. For a given flattened x-ray intensity the required electron beam current at the x-ray target decreases as about the 1.8 power of beam energy; the required electron beam power decreases at about the 0.8 power of beam energy. For example, to produce 500 cGyImin over a 50-cm diameter flattened field at 100-cm SAD, the average electron beam current and power at the x-ray target are approximately 125 yA (microamperes) and 0.75 kW at 6 MV; 17 yA and 0.31 kW at 18 MV-x-ray energy. That is, it takes 7.2 times as much beam current and 2.4 times as much beam power to produce the same dose rate at 6 MV as at 18 MV flattened over a 50-cm diameter at 100-cm SAD. In machines that employ a bend magnet, the actual electron beam current and power within the accelerator guide are about 1.6 or more times these values, because of beam clipping at the input collimator to the bend magnet and at the energy slit within the bend magnet. In machines without a bend magnet, the accelerated beam current is also larger because of its significant low-energy electron content. The gun injects beam into the first cavity of the accelerator guide continuously during the several microsecond pulse.
However, since the rf electric field in each cavity is oscillating at 3 GHz from forward to reverse direction, the electrons are repelled back toward the gun during each 180" of reversed rf electric field. Even during the time when the rfelectric field is forward, electrons are captured over only about 120" of the rf cycle and accelerated. In addition, the input collimator to the accelerator guide may clip some of the injected electron beam. The net result is that the gun emits over three times the accelerated beam current and in bent beam machines, five or more times the current at the x-ray target. Some machine designs require much higher gun currents and accelerated beam currents because of greater clipping of the beam at the collimator and energy slit.
MAJOR SUBSYSTEMS AND COMPONENTS Each manufacturer has its own set of fundamental philosophies and technical approaches to the design of medical electron linacs. There are fundamentally different types of gantry mount, modulator, rf power source, microwave power control, electron gun, accelerator guide structure, energy switching technique, bend magnet, beam distributing system, ionization chamber, safety interlock system, control system, computerization, treatment head, patient table support, treatment beam modifying and shaping devices, and accessories. Table 1-4 summarizes the range of design choices available.
Modulator and high-voltage pulse transformer All linacs employ a modulator, comprising a high-voltage dc (direct current) power supply and pulse modulator. It converts ac (alternating current) mains power to high-voltage pulses, which are applied through a pulse transformer to the cathode of the rf power tube. It may be mounted in a separate enclosure, in the stationary stand that supports the gantry, or in the rotatable gantry itself. The pulse transformer is mounted close to the rf power tube for optimum pulse shape. It may be in the modulator enclosure, but more typically it is in the gantry stand or the gantry. The modulator pulses can be fed through a coaxial cable of many meters length to the pulse transformer. The modulator employs a pulse forming network (PFN) and typically a gas-filled switch tube (thyratron). The PFN comprises a number of capacitors separated by inductors (coils). When the switch tube is triggered it connects the PFN across the pulse transformer primary. The capacitors proceed to discharge their stored energy, but sequentially, because of the time delay produced by each successive inductor. After the pulse is over, the switch tube extinguishes (typically due to a small reverse voltage reflection from the pulse transformer). The capacitors of the PFN then recharge from the high-voltage power supply through a hold-off diode and resonant charging choke to almost twice the power supply voltage, until a De-Q circuit senses that
19
ELEMENTARY DESCRIPTION OF MEDICAL LINACS
TABLE 1 4
. Subsystem and component design alternatives
TABLE 1 4
(Continued)
-
Item
Alternatives
Item
Alternatives
Gantry
Stand supported (internal bearing) Drum type (external bearing) Magnetron "Boot" magnetron Klystron
Control system electronics
Microprocessors Computer (s) Symmetrical Independent X. X and Y Multileaf Manual. Automatic. Dynamic Pedestal (scissors) Extended range (double scissors) Ram (deep floor pit) Carbon fiber. Other X Y translation Eccentric axis rotation Tilt. Pitch
Radio frequency source
Radio frequency power control
Elech-on gun
Accelerator guide
X-ray energy selection Standing wave
Traveling wave Solenoid Bend magnet Nonachromatic Achromatic 270"
45": -45":
+ 112"
X-ray target Primary collimator Beam distribution X-rays Electrons Ionization chambers
Control system electronics
Microwave variator Modulator voltage Klystron drive power Diode Directly heated wire cathode Impregnated indirectly heated cathode Triode Wire mesh grid Focus electrode Modulating anode Standing wave Side coupled Axially coupled, biperiodic, or triperiodic Traveling wave, with or without rf feedback Energy switch in side cavity Nonshorting. Shorting Very short buncher Beam loading. Detuning Beam loading. Detuning Buncher only. Full length (older machines) Radially only. Plus transverse Separated poles Stepped poles Tiltable mid-orbit poles "Slalom" Within vacuum. External Fixed. Round. Square, rotatable Interchangeable for X and E Full flattening filter Scanning, thin flattening filter Scattering foils Scanning Hermetically sealed Open, autocorrection for tempature, pressure, humidity Interchangeable for X and E Discrete components PROMS,e.g.,
Secondary collimator
Wedge filter Patient table support
Patient table top
the correct charge voltage is reached and dumps the remaining stored magnetic energy of the charging choke via a second thyratron into a resistive load. The charging choke and the capacitors of the PFN form a resonant circuit with Q of order 100 and when the De-Q switch fires it spoils this resonant Q, hence " D e - Q (see Appendix C-2 for definition of Q). The gas-filled thyratron may occasionally fire through without being triggered, especially toward the end of its life. For improved machine reliability, it would be desirable to eliminate thyratrons. In the future, other types of modulators may be developed for medical linacs. One type is the magnetic modulator, which achieves pulse compression from a longlow-current pulse to a short-high-current pulse through a series of successively shorter time constant saturable reactors. This pulse can be initiated by solid state switches, thus avoiding use of the thyratron. Solid state modulators have also been built for research linacs, using a number of small PFNs and solid state switches, switching them simultaneously in parallel to the pulse transformer.
Radio frequency power source and radio frequency power control Low-energy linacs (4-8 MV) employ a magnetron rated about 2.5-MW pulse rf power output. The magnetron is an oscillator. Its frequency is determined by the resonant frequency of the cavities machined into its water-cooled cylindrical copper anode, by the phase and amplitude of the rf power reflected to it from the feed to the accelerator guide, and by a motorized tuning plunger in one of the magnetron cavities. During the many seconds of each radiation treatment, the copper inner surfaces of the magnetron anode and of the accelerator guide warm up relative to the cooling water, changing their resonant frequencies but not in synchronism. An automatic frequency control (AFC) circuit senses the resonant frequency of the accelerator guide and drives the motorized tuning plunger in
20
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
the magnetron to maintain approximate synchronism of the magnetron oscillation. A femte device is used to reduce the amplitude of reflected rf power at the magnetron. Ferrite is a magnetic ceramic that has the property of rotating the electromagnetic wave or advancing its phase, such that with appropriate microwave circuitry, reverse flowing rf power can be dissipated. Magnetron rf power output can be controlled by changing the modulator voltage. However, this creates an impedance mismatch, since the magnetron current changes rapidly with small changes in applied voltage. One way to avoid this is to use an electromagnet to establish the magnetron magnetic field and vary the current in the electromagnet as the magnetron voltage is varied. Another way is to run the magnetron at constant values of voltage and current and use a microwave power divider (e.g., ferrite and variable position short) to vary the portion of magnetron power that is fed to the accelerator guide, the remainder being dumped into an rf load. Some medium (10-15 MV) and high-energy (to 25 MV) linacs also employ a magnetron, but of the "Boot" type with an output of about 4.5 MW. The higher output pulse power is achieved by making the cathode-anode structure about four times as long as in the 2.5-MW magnetron, in order to obtain adequate cathode current. The structure length is a much greater fraction of a half-wavelength of the fundamental microwave oscillation frequency,so it is more difficult to suppress oscillation at other frequencies, to maintain uniform current loading over the cathode length, and to suppress arcing. The use of rf feedback with a traveling wave accelerator guide helps to reduce frequency instabilities and hence is used with this type of magnetron in some accelerators. A klystron is used as an rf power source in high-energy linacs that employ a standing waveguide. Since the klystron is an amplifier, an rf driver is required. The rf driver output pulse power needs to be only 100 W. A microwave planar triode can be used as an rfdriver, stabilized in frequency by a temperature controlled low loss rf cavity, resonant in a high-order mode. Alternatively, a solid state driver can be used. It is preferable to run the klystron in the saturated mode. This means there is enough rfdrive power to produce optimum bunching at the klystron output cavity. The output power is then relatively independent of any variations in drive power. At 100-kV klystron voltage, the electrons are at 55 percent of the velocity of light and take about 2 X 10-9 second to travel from the input cavity to the output cavity. This corresponds to about 2000" of rf phase. A 1 percent change in klystron voltage corresponds to an 8" change in phase of the rf output power relative to the phase of the rf input drive power. The amplitude of the rf output power also varies with modulator pulse voltage amplitude. The filling time of the accelerator (to build up the electromagnetic fields) acts as a filter in reducing the effect of these variations on the phase and amplitude of the accelerating electric field. Still, it is important that the modulator provide a voltage pulse that is flat within relatively tight tolerances.
SUMMARY OF ENERGY CONVERSION STEPS In summary, a microwave electron linac comprises equipment to transform electrical energy in a series of steps, from continuous mains power at a few hundred volts ac to successively higher energetic voltages in successively shorter packets. The following approximate values apply to a high-energy medical linac. 1. Alternating current mains power to a high-voltage dc power supply in the modulator: 208 or 380 V, 15 kVA, 50 or 60-Hz input, transformed and rectified to 10-kV dc. 2. Resonant charge of PFN: 10 kV is doubled to 20 kV through a charging choke that resonates with capacitors of the PFN in one-half cycle of 2 X second. 3. Discharge of PFN to klystron: The PFN is switched via a thyratron tube (or solid state devices) and pulse transformer, transferring 100 J (watt seconds) energy stored in the PFN to the klystron cathode at 120 kV in a 7 X s pulse. 4. Radio frequency power from the klystron fills the accelerator guide structure with stored electromagnetic energy of 3 J in the first s and continues to replenish this energy s rf pulse. during a 6 X 5. Radio frequency fields in the accelerator guide transfer energy to the electron beam: At any instant there are about 15 electron bunches each 1 cm long and 10 cm apart in the accelerator guide. Assuming 11 X A at 18 MeV, 0.2 MW is transferred to the beam during 5 X s, 1 J per klystron pulse, 0.7 X J in each of 1.5 X lo4 electron bunches. 6. Conversion of an electron beam to an x-ray beam by the braking action of the electric field of atomic nuclei in the x-ray target: At 18 MeV in a relatively large field of 25 X 25 cm, the x-ray power delivered to a 25-cm thick patient at 500 cGy1min is only 1 W, or less than loL4of the mains power to the modulator.
DESIGN CRITERIA FOR RADIOTHERAPY ACCELERATORS There is a fundamental set of clinical requirements that must be satisfied by any type of radiotherapy accelerator. Each clinical requirement translates into one or more major design criteria and machine performance criteria. Table 1-5 presents a summary of clinical requirements and corresponding machine criteria. The justification for some of these requirements is discussed in the following first few subsections. Some of the challenges that they present to the machine designer are discussed in the subsequent few subsections. Finally, some of the changes in medical linac technology over the past several decades are discussed briefly.
TABLE 1-5 . Some clinical requirements of megavoltage radiotherapy accelerators Item
Major criteria
Precise delivered dose throughout target volume
Precise dimensions of target volume Minimal dose to normal tissue
Wide variety of radiation modalities
Reliability Convenience of patient set up
Moderate time to irradiate, for patient comfort, minimal motion Patient safety
Flatness of fields-all field sizes. Stability of field flatness versus angles of gantry and beam limiting device Stability of penetrative quality Spatial precision of machine and radiation beam Spatial precision of position indicators Depth to maximum dose Penetrative quality Slope of fall-off of electron depth dose Sharpness of dose profile shoulder Width of penumbra at depth Scatter from beam modifiers. Low and high x-ray energies Low to high electron energies Small to large field sizes Rotational therapy Minimal unscheduled down time Set-up time per field Range and ease of equipment motions Height of isocenter above floor Dose rate (with beam modifiers) Mechanical injury avoidance Radiation injury avoidance
TABLE 1-6 . Uncertainty in dose at off-axis point in target volume (95% confidence limit, 2 standard deviations) Uncertainty (%) Calibration Devices (ICRU - 24) Physical constants Standardization of beam at Nat'l Stds. Lab. Calibration of secondary instrument of Reg. Cal. Lab. Calibration of field instrument (optimal model) Calibration of treatment beam (optimal model) Delivery of dose to tissue phantom (optimal model) Calibration root mean square Linac (IEC suggested tolerances) Dose monitoring system: Reproducibility Proportionality (linearity) Dependence on equipment position Stability throughout the day Stability throughout the week Dose monitoring system root mean square Stability of dose due to 2 5 mm tolerance in SSD indicator Stability of dose at depth due to ? 1 percent instability in penetrative quality Stability of flatness with angular position Beam stability root mean square Total linac root mean square Total of calibration and linac root mean square Other (AAPM) Imprecision in treatment planning computation Contributions from organ motion and changes in patient anatomy Total of all contributions (root mean square)
?1.1 20.5 f0.4 21.0 21.7
20.7 2.5
22
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
CLINICAL REQUIREMENTS
Precise delivered dose at depth The goal of curative radiotherapy is to sterilize the cells of the primary tumor without excessive damage to intermingled and surrounding normal tissue. The margin for error in dose can be quite small for some tumors. An analysis of clinical data by Herring et al.35 concludes that since + 10 percent changes in radiation dose can give marked changes in the probability of normal tissue necrosis or of primary tumor recurrence, the therapist needs to be able to deliver a dose distribution to the patient such that the dose at the tumor or at other critical volumes is known to within t 5 Dercent. One authoritative assessmenthas been provided by a committee of the International Commission on Radiation Units and Measurements(ICRU). Section 7.2 of ICRU-2439 states that for certain types of tumors an accuracy of + 5 percent in absorbed dose to the target volume should be the criterion if eradication of the primary tumor is sought. The implication is that this applies throughout the target volume. It further states that at the best level of current practice the uncertainty in calibration of dose in a phantom is + 2.5 percent at a 95 percent confidence limit (2 standard deviations), exclusive of uncertainties related to the treatment machine. Table 1-6 lists additional relevant tolerance values recommended by the International Electrotechnical Commission (IEC)38 for medical electron accelerators. Combining the ICRU and IEC figures randomly for simplicity, the uncertainty in delivered dose to the selected point displaced from the beam axis at depth in the phantom is + 4.2 percent at a 95 percent confidence limit. Uncertainties of 4.2 percent in the treatment planning computation process, as estimated by AAPM,' increase the tolerance to + 6 percent at a 95 percent
confidence limit. Stated differently, in about 9 percent of patients the delivered dose will differ from the isodose treatment plan by more than + 5 percent somewhere in the target volume. In summary, even under optimal conditions the many contributions to dose uncertainty make it difficult to meet the clinical goal of t 5 percent accuracy of dose throughout the target volume in the patient, so high precision of linac performance, ease of precise patient setup and thorough quality assurance procedures arejustified.
precise position, orientation, and size of treatment fields Some organs have a low tolerance for radiation, such as kidneys, spinal cord, lung, liver, and rectum. It is often not possible to avoid irradiating portions of critical organs and still ensure an adequate irradiation target volume to treat the tumor. Low-tolerance organs (and any normal tissue) can tolerate a higher dose without serious injury if the volume exposed to radiation is reduced, as determined by Berg et al.5 and Schultheiss.65 Risk of complication in normal tissue versus dose and portion of organ exposed is summarized56 in Table 1-7. The patient's organs and tumor in the abdominal and pelvic regions may move as much as 1-3 cm relative to bony landmarks according to Chen et al.,l3 due to respiration and changes in patient anatomy during the course of treatment. There are also practical limits to the precision with which patients can be routinely positioned each treatment day. These spatial variations require prescription of treatment fields sufficiently larger than the assumed tumor volume to ensure it is always included in the target volume, which further increases the need for sharp precisely located field
TABLE 1-7 . Cumulative doses of radiation delivered with standard fractionation that have 5 and 50 percent probability of producing fatallsevere morbidity within 5 years
Organ
Injury
Bone marrow
Aplasia, pancytopenia
Liver
Acute and chronic hepatitis
Stomach Intestine
Perforation, ulcer, hermorrhage Ulcer, perforation, hemorrhage
Brain Spinal cord Heart
Infarction, necrosis Infarction, necrosis Pericarditis, pancarditis
Lung
Acute and chronic pneumonitis
Kidney
Acute and chronic pneumonitis
Fetus
Death
TDsls
TDso15
250 3000 2500 1500 4500 4500 5000 5000 4500 4500 7000 3000 1500 1500 2000 200
450 4000 4000 2000 5500 5500 6500 6000 5500 5500 8000 3500 2500 2000 2500 400
Whole or partial organ (field size or length) Whole Segmental Whole Whole (strip) 100 cm 400 cm 100 cm Whole 10 cm 60% 25% 100 cm Whole Whole (strip) Whole Whole
23
DESIGN CRITERIA FOR RADIOTHERAPY ACCELERATORS
I
I
I
I
I
ORGAN SPINAL CORD
-
RECTUM THYROID INTESTINE
-
BONE MARROW
-
I
I
I
I
REFERENCE VOLUME
TOLERANCE' DOSE.cGy
10 CM 100 CM2 100 CM2 LOBE WHOLE 100 CM3 WHOLE LOCALIZED WHOLE
4500 5500 5500 4000 4500 4500 2500 2000 2300
-
'5% PROBABILITY OF INJURY BY 5 YEARS
-
-
-
0 0
FIGURE 1-22
I
I
10
20
I
1
I
I
I
I
I
30 40 50 60 70 80 PERCENT OF ORGAN REFERENCE VOLUME
100
90
. Dose-volume dependence for injury to organ (based on Ref. 54).
edges to avoid injury to neighboring critical organs. The typical dimensional tolerance recommended by the IEC39 is 2 mm for each individual motion and indicator, which results in a cumulative maximum error of 2 4 mm in the position of the x-ray field axis and x-ray field edges relative to the true position in space of the x-ray isocenter. Errors in abutment of adjacent fields also create the potential for localized overdose or underdose, leading to injury to normal tissues or recurrence of the primary tumor. This indicates the need for precise positioning of the edges of the radiation field, not only at the patient's skin but at the depths of critical organs. Since the probability for injury increases rapidly with dose but slowly with the volume exposed (see Fig. 1-22), this problem can be ameliorated by tapering or successively shifting the abutting edges of the radiation fields.
+
What is needed is the capability to vary field sizes over a wide range and still maintain a flat dose distribution with sharp shoulders and small penumbra at the depth of the target volume. For example, x-ray field sizes to 40 X 40 cm at the isocenter with diagonals to about 50 cm are needed for some applications. Above all, the accelerator must be reliable. Unscheduled down-time can play havoc with patient schedules and patient faith in the process and can cause uncertainty in biological dose equivalent. Convenience of patient and beam positioning are essential, since the time to set up patients exceeds the beam on-time by a large factor. The safety for patients and staff are paramount.
140
I
I
Wide variety of radiation modalities Figure 1-23 compares depth dose distribution for parallel opposed 6 and 18-MVx-ray beams through 10- and 30-cm patient thicknesses, corresponding roughly to head and neck tumors versus abdominal and pelvic tumors. For example, for delivery of a given dose to an 8-cm thick treatment volume at mid-depth in a 30-cm thick patient section, the dose to overlying normal tissue is significantly less (12 percent) with an 18-MV x-ray beam. However, for a lOcm thick patient section, the same treatment volume receives a significantly more uniform dose (15 percent) with a 6-MV x-ray beam because of the more rapid build up of dose with depth. Mixed beams of different ratios of high- and low-energy x-ray beams or of x-ray and electron beams provide wide flexibility of treatment plans. Providing this multiplicity of modes in the same machine can result in greater precision and convenience of patient setup.
I
3OCMI THICK
I
I
I
~ ~ I O THICK-] C M
I
0 -20
I
-15
I
I
-10
-5
I I 0 +S CENTIMETERS
I
I
I
+10
+15
+20
RGURE 1-23 . Depth-dose distributionsfor parallel opposed fields (from Ref. 54).
24
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
SOME DESIGN CHALLENGES A fundamental aspect that distinguishes the design of accelerators for radiotherapy from accelerators for other applications is that radiotherapy machines must be people-sized. The patient and the radiation technologist are all-important. Some design challenges related to their needs are discussed in the following subsections.
Compactness Cancer patients are sometimes quite sick and usually need to lie on their backs for radiation treatment. Also, the patient's anatomy shifts markedly from supine to prone positions. In order to irradiate the target volume from different directions without turning the patient over, 360" rotation of the gantry is needed. For convenience in settingup the patient,the isocenter,the point in space around which the equipment rotates, should not be too high above the floor. Adequate space must be provided between the isocenter and the radiation head for radiation technologist access to the patient and for beam modifying accessories. Figure 1-20 shows that this leaves a very limited amount of space for the various components and the radiation shielding in the radiation head, and particularly for the bend magnet or for an inline accelerator. To a significant extent, the design challenge over the years has been to stay within this space while making major advances in the clinical utility of machines and in the treatment beam characteristics, including beam precision and stability and a wide range of x-ray and electron energies.
High-Dose rate with large fields One demanding requirement is to achieve a high-dose rate flattened to the comers of a quite large x-ray field, especially in the low x-ray energy mode, such as for treatment of lymphomas. High-dose rate should be available to limit exposure time, for patient comfort, and to reduce the probability of patient motion during beam on time. A high dose rate beam can be gated on in synchronism with low velocity points in organ motion while maintaining acceptable treatment times. Highdose rate is also needed to allow for absorption in wedge filters and compensatorsand for treatment at extended distances, such as for hemibody irradiation.
Dose precision Another demanding requirement is for precision of delivered dose. Defining gamma as the slope of the curve of tumor control probability (TCP) versus dose at 50 percent TCP, Brahmeg lists clinically observed gammas for a variety of tumor sites and stages. The values of gamma range from 0.4 to 8.0, with a mean value of about 3. This means for a patient with a TCP of 50 percent and gamma of 3, an underdose of 5 percent over the full treatment course will reduce the TCP by about 15 percentage points to 35 percent. Similarly, an overdose of 5 percent may increase the probability of severe damage to normal tissue from 5 percent
to perhaps 9 percent. Also, if the excess dose to a region of 3 percent asymmetric field is in the most critical organ side for the course of therapy, the excess dose can increase the probability of severe injury to that organ.
Beatment beam stability Treatment plans are typically prepared from a set of isodose contours measured or calculated for a single gantry angle and single beam limiting device angle. Although difficult, stable unflatness of fields could be accounted for in the treatment planning process to achieve the desired uniformity of dose over the target volume. A really insidious contribution to error in delivered dose distribution is then the potential instability of x-ray field flatness with rotation of the gantry and beam limiting device, due to instabilityof position and orientation of the electron beam at the x-ray target relative to the axis of the flatteningfilter. The problem is illustratedin Fig. 1-20for an 18-MVx-raybeam. A point P in the field 10 cm from the isocenter intercepts a ray B from the point source centered at the x-ray target, but point Pintercepts ray A from a point displaced by 10 mm at the x-ray target. The attenuation by the flattening filter is 3:l (attenuating from 300 to 100percent intensity) for ray B but 6:l (attenuating from 300 to 50 percent intensity) for ray A. Thus, a 10-mmdisplacement of the electron beam on the x-ray target causes 100 percent field asymmetry, down to a 50 percent dose on one side of the field, up to a 150 percent on the other side. However, a simultaneous tilt of the electron beam by 100 milliradians realigns the x-ray lobe with the peak of the flattening filter (attenuatingfrom 600 to loopercent intensity) alongray A, canceling the asymmetry at the point P. Scaling the above exaggerated illustration linearly, in order to maintain field symmetry within 1%, the mean position and the mean angle of the electron beam at the x-ray target must be individually maintainedwithin about -t 0.1 rnm and 1milliradiansfor an 18-MV x-ray beam. It is important that the bend magnet system be doubly achromatic so that the angle and displacement from axis of each ray in the electron beam out of the accelerator guide be reproduced in the electron beam leaving the magnet independent of the energy of each ray. This permits unambiguous feedback from radiation beam monitors to maintain precisely the mean position and angle of the beam at the x-ray target independent of any changes in mean energy of the spectrum of energies transmitted by the energy slits in the magnet. The magnet system should be achromatic both for rays displaced from the central ray in the plane of bend (radial plane) and for rays displaced transversely to this plane (transverse direction). Instabilities in main beam energy can produce symmetrical instabilities in field flatness, especially at high x-ray ener1 gies resulting from variation of the x-ray lobe width. A percent error in beam energy at 18 MeV will produce about a 1.7 percent error in dose near the periphery of a 40 X 40-cm field relative to the dose on the axis. Such a shift in mean energy could occur due to changes in the shape of the energy spectrum transmitted by the energy slits in the bend magnet, which typi-
+
+
+
25
DESIGN CRITERIA FOR RADIOTHERAPY ACCELERATORS
cally transmits a 6 percent energy bin. (Some manufacturer's machines transmit a 10 percent or even larger energy bin). A full-field ionization chamber is often used for safety and for high signal-to-noise (S/N) ratio. However, it will respond primarily to dose away from the axis as pointed out by Sutherland.70 Since calibration with a field instrument is routinely done on the beam axis, this alone may not be enough and additional means of energy regulation may be needed. For example, by operating in a region where the x-ray dose rate is a sufficiently fast function of beam energy, feedback from the dose rate monitoring system to the microwave power source can be used to regulate mean energy of the beam at the x-ray target.
Uniform x-ray treatment beams, with minimal contamination The x-ray beam is emitted from the target in a forward lobe that is clipped by a primary collimator, flattened by a conical flattening filter, monitored by a multiple section ionization chamber, clipped further by movable jaws and multileaf collimator, and shaped by accessories mounted on optically transparent trays (see Fig 1-20). Each of these items and the intervening air contribute lower energy scattered photons and electrons that increase the dose to the sensitive vascular layer of the patient's skin and modify the shoulder and penumbral regions of the dose profile. The lateral transport of secondary electrons from photon interactions in the phantom increases as the x-ray energy is increased, thereby softening the shoulder and increasing the penumbral width of the dose profile. For x-ray energies of 15 MV and above, Monte Carlo calculations by Mohan et a1.52 show that the contribution to shoulder softening and penumbra resulting from secondary electrons produced in the phantom exceeds the total contribution from machine sources. Nevertheless, for those aspects within the control of the manufacturer the total of effects can be minimized by the proper choice of materials in the beam and by proper orientation of the primary collimator and jaw faces of the beam limiting device.
materials and their location and shape to limit production of lower energy and obliquely scattered electrons and production of x-rays. For example, thin light weight low atomic number material can be used for the several layers of the dose monitor ionization chambers. Fig. 1-25 shows that the depth dose distribution for electrons from a linac is essentially the same as for the very narrow energy spread electrons from a microtron.
Energy stability The IEC39 suggested tolerance on stability of depth dose in the electron mode corresponds to -+ 1 percent energy stability at energies above approximately 10 MeV. That is, the deviation of the mean of the energy spectrum transmitted by the energy slit should be limited to 1 percent. This can be facilitated by accelerating a broad fairly uniform energy spectrum such as 20 percent width and selecting out only a portion of this spectrum by using a narrow energy slit such as 6 percent. Electron beam current can be wasted in the electron mode because for a given dose rate the required beam current to be delivered to the electron window is less than 1 percent of the required beam current at the x-ray target in the x-ray mode. A narrow energy slit for the electron mode can limit the dose rate in the x-ray mode if the same slit is used and if the accelerated beam energy spread is not sufficiently narrow. If a high dose rate is to be achieved at low x-ray energy fully flattened to the comers of a large x-ray field, the percentage beam current transmission through the bend magnet system to the x-ray target must be relatively high (see Fig. 1-24). To pass through a narrow energy slit this requires that the gun inject a beam with low transverse emittance (displacement and angular dispersion of electron rays within the beam) into the accelerator guide and that electrons ride the accelerating wave in such a way as to avoid instabilities and increases in energy spread. Also, high electron beam transmission through the electron beam collimator and energy slit permits thinner, lighter weight shielding in the radiation head.
+
I
Uniform electron treatment beams with minimal contamination The electron beam can be spread by multiple scattering foils into a uniform treatment beam that passes through an ionization chamber and is shaped by an applicator insert. These items and the intervening air contribute x-rays that produce dose outside (beyond) the intended treatment volume and contribute lower energy scattered electrons that modify the depth dose distribution. These scattered electrons increase the skin dose and reduce the clinically useful depth of the 90 or 80 percent depth dose for a given mode energy. The obliquely scattered electrons also degrade the transverse distribution of dose, resulting in poor penumbra when a portion of the applicator cannot be close to the patient's skin, such as in the neck region when the chin interferes. These effects can be minimized by proper choice of
I
I
I
I
I
I
I
I
I
-
6X MODE -
6% SLIT
18X MODE
4!--
6% SLIT
4 k
I
-
i i
I
4
I
II II
I
0 0
2
4
6
8
10
12
14
16
18
20
22
ENERGY (MEV)
FIGURE 1-24 . Energy spectra of electron beam in microwave switched 6 and 18-MV x-ray modes (from Ref. 54).
26
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
frequency power feedback was employed over the accelerator guide to reduce variations of beam energy due to magnetron frequency instabilities. The vacuum system employed an oil diffusion pump, which remained vertical as the gantry rotated. Because the in-line accelerator guide was so long, the gantry could be rotated through only 120°, from 15"beyond the vertical to 15' below the horizontal, or 30" below the horizontal by retracting a panel in the floor of the treatment room. Because of limited microwave power, the typical dose rate was only 100 cGyImin with beam flattening over most of only a 20 X 20 cm maximum field size at a 100-cm SAD. The average treatment load was 30 patients per day. The mean time between failure of some of the components was magnetrons, 1Omonths;high-voltage rectifiers, 4 months; ignitrons, 1 month; other vacuum tubes, 1 week; electron gun filaments, 3 months. Table 1-8 summarizes some of the changes in medical linac technology from early-to-modem machines and the resulting improvements in clinical performance. In general, these relate to improved accelerator guide and magnet systems and beam modifying systems to provide wide ranges of beamenergy, dose rate, field size, and operating modes with improved radiation beam characteristics in compact machines of high reliability.
10
-
0 0
I
I
2
4
,
SUMMARY ACCELERATOR MAJOR SUBSYSTEMS I
6 8 CENTIMETERS DEPTH
10
'2
14
FIGURE 1-25 . Comparison of electron depth dose curves for linac and microtron (from Ref. 54).
Initial seconds The beam energy, dose distribution, and dose calibration should be within tight tolerances even during the initial seconds of each portal treatment. In the future, conformal therapy may call for say 10-port treatments of 5 s each. A 10 percent asymmetry through just the first second could add 2 percentage points to the steady state value of asymmetry. Assuming a tumor dose of 200 cGy, a 0.1 cGy round-off error or depth dose error in the 20 cGy to each of 10 portals could add another 0.5 percent dose error. These errors could become significant if they add systematically. A similar rationale applies with hyperfractionation where there are more beam initiations in a course of treatment. CHANGES IN TECHNOLOGY FROM EARLY-TO-MODERN MACHINES The first single gantry isocentric medical electron accelerator design was developed in England in the early 1950s and was described by Miller.51 The 4-MV x-ray beam was emitted in line with a 100-cm long traveling-wave accelerator guide. Radio
Essentially all modem radiotherapy machines employ an electron linac mounted in an isocentric gantry. Accelerators are loosely classified by their x-ray and electron energies as shown in the following table:
Machine class
X-ray mode
0 Low energy
Medium energy High energy Dual energy
4 or 6 8 o r 10 15118/20/22/25 4 or 6 and 10115/18/20/22/25
Electron mode (MeV) None 416 to 9/12' 416 to 12/15 416 to 18/25 416 to 18/25
' I means or
All low-energy machines employ a magnetron, which produces a peak rf power of about 2.5 MW, with 2.5-kW average rf power. Medium, high, and dual energy machines employ a larger magnetron or a klystron. The rfpower is pulsed on by a modulator for a few microseconds at intervals of a few milliseconds. With each pulse, the accelerator delivers a dose at the isocenter of typically 0.01 to 0.03 cGy. The modulator contains a high-voltage power supply, charging voltage regulator (De-Q circuit), PFN, and thyratron switch tube. It pulses the magnetron or klystron through a pulse transformer, which
27
DESIGN CRITERIA FOR RADIOTHERAPY ACCELERATORS
TABLE 1-8
. Some changes in medical linac technology from 1950s-1990s
Item Accelerator guide type
Early naveling wave
Modem Standing wave
4
(37-47)
12-18 (86-1 12)
Nonachromatic
Achromatic
X-ray field size
Modest
Large
X-ray dose rate centrigray per minute (radslminute) X-ray energies, MeV (number of modes)
100-200
Mev per meter of guide (shunt impedance, megohmslmeter) Bend magnet
4-6 (1)
None or low
Low to high
Fair
Excellent
Months
Years
Cleanliness
Oil pumps
Ion pumps brazed guide
Electronics
Tubes and relays
Solid state modular
Electron energies Isodose distributions and their stability Microwave tube life
steps the voltage up from about 10 kV to about 50 kV for magnetrons, to about 120 kV for klystrons. An automatic frequency control system (AFC) senses the resonant frequency of the accelerator guide and adjusts the frequency of the rf source via a tuning plunger in the magnetron or via the rf driver to the klystron. The dosimetry system servos beam pulse rate via the gun or rf source modulator to control dose rate. In low-energy machines without a bend magnet, x-ray energy is measured indirectly as the ratio of dose rate to the electron beam current impinging on the x-ray target. In machines with a bend magnet, the energy is defined by the energy slit and bend magnet current. The accelerated beam energy can be maintained by feedback from the dose rate monitor to control the amplitude of rf power.
ONE FUTURE DIRECTION OF EQUIPMENT DEVELOPMENT IN RADIATION THERAPY The first revolution in the treatment of cancerwas the invention in the 1950's of machines to produce penetrating Megavoltage radiation beams (cobalt, various types of electron accelerators).
Result Doubled guide efficiency
Shorter guide, simpler, more compact machine. 360" gantry rotation Stable treatment fields Full mantle at isocenter Short exposure, even with wedge filters Optimal for thin and thick sections of patient Full useful penetration Protection of normal tissue, dose precision Machine uptime lower cost Freedom from arcing High energy gradients Reliability ease of service
Because of its superior beam characteristics (penetration, precision, versatility, dose rate), the electron linear accelerator (linac) has become the machine of choice among radiation oncologists. In the industrialized world, over 75% of radiation machines are linacs. The second revolution (in the 70s and 80s) was based on advances in computer power and the invention of the Computerized Tomography scanner (CT) and the Magnetic Resonance Imager (MRI). These developments have made radiation oncology a precise and predictive modality. Tumors and surrounding normal tissues could now be localized and characterized with accuracy. However, because of various limitations, treatment planning was still limited to 2-dimensional visualization and 2, 3 or 4 co-planar treatment fields. Even so, relative 5-year survival rates7 in the United States averaged over all tumor sites in which patients (white) were treated with curative intent have increased from 39% in the early 1960's to 53% in the mid-1980's and would be close to 60% if lung cancer were excluded. All cancer sites amenable to radiation therapy treatment have shared in this improvement. The third revolution, 3-dimensional conformal radiation therapy (3-D Conformal RT) is now in progress. Advances in the ratio of computational power to computer cost are now making it economically practicable. The data from up to 60
28
CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
CT scans (slices) and from MRI data are reformatted by high speed graphic computer workstation to provide 3-D displays of the patient's anatomical structure. Accelerator beam dose distributions are then computed and superimposed on this anatomical display for beams from many directions in a search for the optimal treatment plan for that individual patient. By shaping the outlines of the treatment beam and aiming it at the tumor from many directions, the high dose region in the 3-D conformal treatment plan can be concentrated in the target volume (tumor plus appropriate safety margin) while avoiding excess dose to radiation sensitive surrounding normal tissue. An example is treatment of the prostate and associated seminal vesicles while minimizing dose to the nearby rectum and bladder. 3-D visualization from many directions is crucial to ensuring for example that a frond of the tumor is not missed or under-dosed (which would lead to cancer recurrence) but that the high dose region is not made unnecessarily large, with consequent damage to normal tissue. Radiation oncologists look for at least 10 percentage points improvement in cure rates for several sites and 20% improvement has been predicted from computer modelling for the nasopharynx. Conformal radiation therapy was originally proposed by S. Takahashi (see Chapter 2) in Japan in 1975. Japanese industry built several multileaf collimators (MLC) and Japanese radiation oncologists did treat patients with the technique but the development of computational power and of MLC's was less than desired for precision 3-D Conformal RT. In conventional radiation therapy, typically two or three or four treatment fields at different angles (portals) are used. Because the operator must go into the treatment room to change field shaping devices for each portal, it becomes excessively time consuming to use more than about four portals. In 3-D conformal RT optimal treatment may require 10 or 12 or more
FTGURE 1-26
portals. This becomes practical through use of a precision MLC and computerized remote control of movements of the accelerator, the patient table and the MLC. Thus, the operator needs to enter the treatment room only at the beginning and end of the many-portal treatment session. Even without 3-D Conformal RT, the advent of the precision MLC will change the face of radiation therapy. For many tumor sites, the extra effort to do the treatment planning for 3-D Conformal RT may not be justified. For these sites, conventional radiation therapy can be used with two to four portals. Because of the remote controlled MLC, as well as computerized linac motions, the operator's activities are markedly reduced, with consequent major saving in time and cost of patient treatment. Fig. 1-26 shows one type of precision MLC. It is designed as an accessory so that it can be mounted on an existing accelerator radiation head below the collimator jaws. It employs a multiplicity of heavy metal bar leaves driven relative to two sub-frames which are driven relative to two jaws of the rectangular field collimator by small electric motors under computer control. The bars project to 1 cm steps at 100 cm SAD and can travel 16 cm beyond field axis. This extended travel capability with limited length leaves is a result of the use of the two travelling sub-frames. This extended travel can be used not only for offset fields but also for dynamic wedge and dynamic compensator shaping of treatment field dose distributions. The superiority of 3-D Conformal RT over the best conventional RT must be proven by clinical trials. It must be demonstrated that higher dose to the tumor results in significantly greater cure rates without increased damage to normal tissue. For each tumor site, 1000 patients are required to demonstrate with 90% confidence that a 10% improvement in cure rate has been achieved. This takes many years. The most likely sites for such initial clinical trials are cancers of the
. Multileaf collimator (MLC) mounted on radiation head below collimator jaws. a) Axial view, h ) With patient (from Varian).
REFERENCES
prostate, head and neck, and brain. Future trial candidates may include gynecological and other sites. There is mounting clinical evidence that increased tumor dose results in increased cure rate. For example, Sandlefi3 reports that the Patterns of Care Study shows 5-year local failure free rates for Stage C prostate cancer of 64% with 6000 to 6499 cGy dose, 81% above 7000 cGy dose. This represents 17% of 25000 Stage C prostate cancer patients per year, curing over 4000 more patients per year. The challenge is to deliver over 7000 cGy to the entire cancerous region without increased harm to normal tissue. With conventional RT, the above Patterns of Care Study showed a doubling of serious complications from 3% to 6%. The cost of failure to cure far exceeds the extra cost of planning and delivering 3-D Conformal RT. Extremely expensive patient care is required for patients that fail treatment. To facilitate delivery of precision radiotherapy such as 3-D Conformal RT, in addition to MLC's, manufacturers are providing electronic portal imaging systems to confirm proper aiming of the beam relative to the patient, RecordNerify systems to confirm correct machine set-up parameters, C T option on radiation therapy simulators to aid in precise positioning of the patient, and networks to integrate all this information. In the future, all major manufacturers of radiation therapy equipment will need to be able to supply systems for 3-D Conformal Treatment Planning, including advanced computer graphics workstations and associated software. It will take a few more years to make 3-D Conformal RT a reality. The time to do the 3-D Conformal treatment planning must be reduced to economical levels. Advances in software are contributing to this. Training dosimetrists to outline organs in the many CT scan images is relieving the demand for the scarcest resource, namely radiation oncologist's time. The safety of 3-D Conformal RT must be proven. Closing the high dose volume down around the tumor 3-D shape (plus margin) and then increasing the dose has its risks. The tumor position must be set-up daily by the operators and immobilized (or tracked dynamically), which is quite difficult for some tumor sites and patients. There is potentially a higher possibility of error in patient set-up, consequently missing a portion of the tumor some of the time. Hence, much greater operator care will be required. This implies the need for extensive operator training and the need for set-up, immobilization, and tumor position tracking devices to assist the operator. Stereotactic devices for different body regions, feedback from on-line electronic portal imager to robotic controls of patient support, physiologic gating of high dose rate treatment beam synchronously with organ motion, etc., may be appropriate operator aids. What is being learned in the development of 3-D Conformal RT procedures is so valuable and logical that there will be a strong tendency for both research and nonresearch hospitals to equip themselves and start applying the technique in at least a limited way as well as in so-called conventional radiation therapy.
29
REFERENCES 1. AAPM Report No. 13: Physical aspects of quality assurance in radiation therapy. New York, AAPM, 1984. 2. Austin NA: Electronic weapon against cancer. Electronics April 6: 88-92,1964. 3. Avery RT: Electronic accelerator with specific deflecting magnet structures and x-ray target. U.S. Patent 3,360,647 Aug. 1, 1960, issued Dec. 26, 1967. 4. Becker GE, DA Caswell: Operation of a six-MeV linear electron accelerator. Rev Sci Instr 22: 402-405, 1951. 5. Berg NO, M Lindgren: Relation between field size and tolerance of rabbit's brain to roentgen irradiation (200 kV) via a slitshaped field. Acta Radiol (Therapy) 1: 147-168, 1963. 6. Boot HAH, JT Randall: Historical notes on the cavity magnetron. IEEE Trans. Electron Devices, Vol. Ed-23 726729,1976. 7. Ca-A Cancer Journal for Clinicians; 42, (No. l), pp 19-43, 1992. 8. Brahme A: Dosimetric precision requirements in radiation therapy.Acta Radiol 23: 373-391, 1984. 9. Brannen E, H Froelich: JAppl Phys 32: pp 1179-1 180, 1961. 10. Brown KL, WG Turnbull, PT Jones: Stepped gap achromatic bending magnet. U.S. Patent 4,425,506, filed November 19, 1981, issued Jan. 10 1984. 11. Catterall M, DK Bewley: Radiotherapy and the physics of radiation beams, in RJ Post (Ed): Medical research council cyclotron unit silver jubilee book. MRC Cyclotron Unit Hammersmith Hospital, London 1980; pp 57-66. 1la. Caterall M, DKBewley:Fast neutrons in the treatment of cancer. Academic Press, London and Grune & Stratton, New York; 1-393,1979. 12. Charlton EE, WF Westendorp, LE Dempster, and G Hotaling: A new million-volt x-ray outfit. J Appl Phys 10: 374-385, 1939. 13. Chen GTY, M Kessler, WM Saunders: Organ movement: Implications for CT based treatment planning, Abstract No.6 Med Phys 11: p 393, 1986. 14. Chodorow M, EL Ginzton, WW Hansen, RL Khyl, RB Neal, WKH Panofsky: Stanford high-energy linear electron accelerator (Mark 111). Rev Sci Instr 26: 134-204, 1955. 15. Chodorow M, EL Ginzton, IR Nielsen, S Sonkin: Design and performance of a high-power pulsed klystron. Proc., IRE 41:1584-1602,1953. 16. Coia LR, GE Hanks, K Martz, A Steinfield,JJDiamond, S Kramer: Pmctice patterns of palliativecare for the United States 1984-1985. Int J Radiat Oncol Biol Phys 14:lpp 261-1269,1988. 17. Collins GB: Microwave Magnetrons. M.I.T. Radiation Laboratory Series, New York, McGraw-Hill, 1948; vol6,7-8. 18. Cutler CC: Bell Telephone Laboratories Report MN-44-160218, 1944. 19. Day MJ, FT Farmer: The 4 MeV linear accelerator at Newcastle upon Tyne. Br J Radiol 31: 669482,1958. 20. Diamond JJ, GE Hanks, S Kramer: The structure of radiation oncology practices in the continental United States. Int J Radi Oncol Biol Phys 14: 547-548,1988. 21. Enge HA: Achromatic magnet mirror for ion beams. Rev Sci Instr 34385-389, 1963. 22. Fry DW: The Linear Electron Accelerator. Philips Tech Rev 14: 1-32, 1952. Also published in Delngenieur 64:0.51, 1952.
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CHAPTER 1. THE MEDICAL ELECTRON ACCELERATOR
23. Fry DW, RBR-S-Harvie, LB Mullett, W Walkinshaw: Travelling-wave linear accelerator for electrons. Nature (Sept. 13) 160 (NO.4063): 351-353,1947. 24. Gabriel S, R Stanton, DA Berkowitz: Linear accelerator selection for the one machine department. Preprint of paper presented November 1987 at RSNA. (Robert Stanton, Ph. D., Cooper HospitaVUniversity Medical Center, Camden, NJ.) 25. Ginzton EL: The $100 idea. IEEE Spectrum 12 (No.2): 30-39, 1975. 26. Ginzton EL, WW Hansen, WR Kennedy: A linear electron accelerator. Rev Sci Instr 19: 89-108, 1948. 27. Ginzton EL, KB Mallory, HS Kaplan: The Stanford medical linear accelerator; I. Design and development. Stanford Med Bull. 15:123-140, 1957. 28. Ginzton EL, CS Nunan: History of microwave electron linear accelerators for radiotherapy. Int J Radi Oncol Biol Phys 11: 205-216,1985. 29. Goer DA: Radiation therapy treatment: The role of treatment aids and accessories. IEEE Trans Nucl Sci NS-30: 1784-1787, 1983. 30. Haimson J, CJ Karzmark: A new design 6 MeV linear accelerator system for supervoltage radiotherapy. Br J Radiol 36: 650-659, 1963. 3 1. Hall LD, JC Helmer, RL Jepsen: Electrical vacuum pump apparatus and method. U.S. Patent 2,993,638, filed July 24, 1957 issued July 25, 1961. 32. Hansen WW: High efficiency resonant circuit. U.S. Patent 2,190,712, filed July 27, 1936, issued Feb. 20, 1940. 33. Hansen WW: A type of electrical resonator. J Appl Phys 9: 654-663,1938. 34. Henderson WJ, H Le Caine, R Montalbetti: Natrrre (London) 162: 699,1948. 35. Herring DF, DMJ Compton: The degree of precision required in the radiation dose delivered in cancer radiotherapy. Br J Radiol Special Report Series No. 5: 51-58, 1981. 36. Howard-Flanders P: The development of the linear accelerator as a clinical instrument. Acta Radiol Srcppl 1 16: 649-655, 1954. 37. Howard-Flanders P, GR Newberry: The gantry type of mounting for high voltage x-ray therapy equipment. Br J Rodiol23: 355-357, 1950. 38. International Commission on Radiation Units and Measurements ICRU Report 24: Determination of absorbed dose in a patient irradiated by beams of X or garnrna rays in radiotherapy proceditres 7910 Woodmont Ave., Washington, DC, ICRU, Sept. 15,1976. 39. International Electrotechnical Commission Technical Report 977: Medical electrical eqitipment, medical electron accelerators in the range IMe V to 5OMe V: Griidelirtesforficrzctiortal perfonnancecharacteristics. Geneva, Switzerland, IEC, 1989. 40. Kapitza, SP, VN Melekhin: Tlte microtron. London, Harwood Academic Publ, 1978. 41. Kaplan HS, MA Bagshaw: The Stanford medical linear accelerator; 111. Application to clinical problems of radiation therapy. StanfordMed Bull 15:141-151, 1957. 42. Karzmark CJ: Advances in linear accelerator design for radiotherapy. Med Phys 11:105-128, 1984. 43. Karzmark CJ, NC Pering: Electron linear accelerators for radiation therapy: History, principles and contemporary developments. Phys Med Biol 18:321-354, 1973.
44. Kerst DW: The accelerator of electrons by magnetic induction. Phys Rev 60: 47-53, 1941. 45. Knapp EA, BC Knapp, JM Potter: Standing wave high energy linear accelerator structures. Rev Sci Instr 39: 979991, 1968. 46. Laughlin JS, R Mohan, GJ Kutcher: Choice of optimum megavoltage for accelerators for photon beam treatment. Int J Radi Oncol Biol Phys 12: 155 1-1557, 1986. 47. Laughlin JS: Development of the technology of radiation therapy. Radiographics 9: 1245-1266, 1989. 48. Loevinger R, CJ Karzmark, M Weissbluth: Radiation therapy with high-energy electrons-Part I: Physical considerations: 10 to 60 MeV. Radiology 77: 9 6 9 2 7 , 1 9 6 1 . 49. Meredith WJ: Some aspects of supervoltage radiation therapy. Br J Radiol 79: 57-63, 1958. 50. Miller CW: An 8 MeV linear accelerator for x-ray therapy. Proc IEE 101: 207-222, July 1954 and 102:500, Aug. 1955. 51. Miller CW: Linear accelerators for x-ray therapy. Eighth International Congress of Radiology, Mexico City, 1956, also published by Metropolitan-Vickers Co. Ltd., (now Associated Electrical Industries), Trafford Park, Manchester, U.K., 1956. 52. Mohan R, C-S Chui: Validity of the concept of separating primary and scatter dose. Med Phys 12: 726-730, 1985. 53. Mullett LB, BG Loach: Experimental work on corrugated waveguides and associated components for linear electron accelerators. Proc Phys Soc 61: 271-285, 1948. 54. Nunan CS: Design and performance criteria for medical electron r B 1011I: 881-887, 1985. accelerators. Nrrcl h ~ s t Meth 55. Nunan CS: Microwave electron accelerators, in EC Okress (ed): Microwave Power Engineering. New York, Academic Press, 1968,296-325. 56. Perez CA, LW Brady: Principle and practice of radiation oncology Philadelphia, JP Liddicoat, 1987. 57. Quastler H, GD Adams, GM Almy, SM Dancoff, A 0 Hanson, DW Kerst, HW Koch, LH Lanzl, JS Laughlin, DE Riesen, CS Robinson Jr., VT Austin, TG Kerley, EF Lanzl, GY McClure, EA Thompson, LS Skaggs: Techniques for application of the betatron to medical therapy. An1 J Roentgen01 Radium Ther 61 : 2-36, 1949. 58. Rand RE: Recircitlating Electron Accelerators. New York, Harwood Academic Publ, 1984. 59. Readhead PA, H Le Caine, WJ Henderson: Cana J Res A28: 73, 1950. 60. Rosander S, M Sedlacek, D Wernholm: The 50 MeV racetrack microtron at the Royal Institute of Technology: Nucl Instr Meth Stockl~olm204: 1-20, 1982. 61. R-Sherbsie-Harvie RB: Traveling wave linear accelerators. Proc Pl~ysSOC61 :255-270, 1948. 62. Sable M, WG Gunn, D Penning, A Gardner: Performance of a new 4 MeV standing wave linear accelerator. Radiology 97: 169-174, 1970. 63. Sandler HM, DL McShan, AS Lichter: Potential improvement in the results of irradiation for prostate carcinoma using improved dose distribution. Int J. Radi Oncol Biol Phys 22: 361-367,1992 64. Schriber SO, EA Heighway: IEEE Trans NS-22 No. 3: 1060, 1975. 65. Schultheiss TE, CG Orton, RA Peck: Models in radiotherapy: volume effects. Med Pl~ys10:410-415, 1983.
REFERENCES
66. Schulz MD: The supervoltage story. Am J Roentgen01 Rad Therapy Nucl Med 124541-559,1975. 67. Schwinger J: Phys Rev 75:1912, 1949, (and originally quoted by Schiff LI: Rev. Sci. Instr. 17:6, 1946). 68. Skaggs LS, GM Almy, DW Kerst, LH Lanzl, EM Uhlmann: Development of the betatron for electron therapy. Radiology 50: 167-173, February 1948. 69. Skaggs LS, LH Lanzl, RT Avery: A New Approach to Electron Therapy, Second Infer Conf Peaceful Uses Atomic Energy 26:312-316, 1958. 70. Sutherland WH: Dose monitoring methods in medical linear accelerators. Br J Radio1 42: 864, 1969. 71. Tanabe E, G Meddaugh: Variable energy standing wave linear accelerator structure; in Proceedings 1981 Linear Accelerator Conference Santa Fe, NM, Oct. 1981. 72. Taylor T, G Van Dyk, LW Funk, RM Hutcheon, SO Schriber: Therac 25: A new medical accelerator concept. IEEE Trans NS-30 N0.2:1768-1771, 1983. 73. Uhlmann EM, CL Hsieh, CL Lootens: The linear accelerator as a source of fast electrons for cancer therapy. Radiology 66:859869,1956.
31
74. Varian D: The Inventor and the Pilot. Palo Alto, Pacific Book Publishers, 1983. 75. Varian RH, SF Varian: A high frequency oscillator and amplifier. JAppl Phys 10:321-327, 1939. 76. Veksler VI: Proc USSRAcad Sci 43:346, 1944. J Phys USSR 9: 153,1945. 77. Walkinshaw W: Comparison of efficiency of resonant and traveling wave linear accelerators. Proc Phys Soc 61: 246-254, 1948. 78. Weissbluth M, CJ Karzmark, RE Steele: The Stanford medical linear accelerator: 11. Installation and physical measurements. Radiology 72: 242-253, 1959. 78a. Wernholm 0 : The 1200 MeV synchrotron at the University of Lund. Arkiv For Fysik 26: 527-573, 1964. 79. Woodyard JR: A comparison of the high frequency accelerator and betatron as a source of high energy electrons: (Abstract). Phys Rev 69:50,1946 (Presented at November 30-December 1, 1945 Meeting of the American Physical Society). 80. Zatz LM, CF von Essen, HS Kaplan: Radiation therapy with high-energy electrons-Part 11: Clinical experience: 10 to 40 MeV. Radiology 77: 928-939, 1961.
C H A P T E R
Radiotherapy Modalities
A number of different types of machine produced radiation beams are used for radiotherapy. These include x rays, electrons and other particles such as neutrons, protons, and pi minus mesons, as well as heavier nuclei such as carbon, helium, neon, and silicon ions. Therapeutic x-rays vary in energy from about 10 keV to 50 MeV. Such beams are not monoenergetic; they contain a spectrum of photon energies extending from zero to a maximum that corresponds to the energy of the electron producing it. The mean value of the electron beam energy is used to define the x-ray energy. To distinguish x ray from electron energy, the term megavolt (MV) is frequently used for x rays, whereas, million electron volts (MeV) is always used for electrons. Megavoltage x rays, extending from about 1 to 50 MeV (typically in the range of 4 to 25 MeV), are the most widely used radiotherapy modality and are the focus of this book. Therapeutic radiation beams are described by their central axis percent depth absorbed dose curves, by isodose distributions, and by dose profiles. The absorbed dose of any ionizing radiation is defined as the energy imparted to matter by ionizing radiation per unit mass of the irradiated material at the point of interest. The unit of absorbed dose is the gray (Gy), or in non - SI units the rad, where 1 Gy = 1 Jlkg and 1 rad = 0.01 Gy = 1 cGy. (When "dose" and "depth dose" are used, "absorbed dose" and "depth absorbed dose" are intended.) Depth dose curves portray the relative energy deposition as a function of depth on the axis of a normally incident beam in some standard medium such as water, (see Figure 2-1). Isodose distributions are most often twodimensional (2-D) curves of constant dose in water that are normalized to 100 percent at the dose maximum point on the central axis. They are plotted in planes containing the central axis of the beam or in planes parallel to the surface. Such isodose curves are usually plotted in multiples of 10% dose and are provided for the field sizes in use (see Fig. 2-2). In one alternate representation, dose profiles are obtained from transverse plots passing through the central axis at specified depths. Both depth dose and isodose curves are normalized to 100 percent at the dose maximum point (d,,,) on the field axis in water, and the dose rate is defined at this point.
Comparisons between x-ray beams of different energies are often made for 10 X 10 cm beams at 10-cm depth.
ORTHOVOLTAGE X-RAY THERAPY Orthovoltage x rays extend in energy from approximately 100 to 400 kV. The efficiency of x-ray production is low at orthovoltage energies, the x-ray dose rate is low, and shorter source to skin distances (SSD) are employed to compensate for the low output. Such beams deposit maximum dose on or within a few millimeters of the skin surface and attenuate rapidly with increasing depth, in part, because the inverse square reduction with distance varies rapidly for short SSDs. These characteristics often limit the therapeutic dose that can be delivered to deeply lying tumors because of dose limitations of overlying normal tissue. Arepresentative orthovoltage depth dose curve is shown in Figure 2-1 together with two megavoltage curves. Figure 2-2 illustrates isodose curves for the orthovoltage and 6-MV x-ray
Depth In Water (cm)
FIGURE 2-1 . Central axis depth dose curves for 10 x IO-cm x-ray beams in water. The megavoltage curves have admx depth of overlying water. (a) ZOOkV, 50cm SSD, HVL 2.5 mm Cu. (b)6MV, 100 cm SAD. (c) 24 MV. 100 cm SAD.
2
34
CHAPTER 2. RADIOTHERAPY MODALITIES
FIGURE 2-2 . Isodose curves for (a) the 200-kV and (b)6-MV x-ray beams described in Figure 2-1. The 6-MV curves have been shaped by a beamflattening filter optimized for a 15 X 15-cm beam.
beams defined in Figure 2-1. The 200-kV depth dose (Fig 2-la) starts at 100% at the phantom surface and attenuates rapidly ) more with depth. The associated isodose curves (Fig. 2 . 2 ~are curved than the megavoltage curves of Figure 2-2b, due to scatter at large angles being much more probable at low than at high primary photon energies. The 10 and 20 percent orthovoltage isodose curves, lying outside the geometrical edges of the beam, arise primarily from photon scatter radiation. Orthovoltage beams are identified by a half-value layer (HVL) defined as the depth in aluminum or copper - that reduces the transmitted intensity to one-half of that incident under standard conditions. A depth dose of about 35 percent at 10 cm depth in water, with the dose maximum lying close to the surface, is representative for a 200-kV (HVL = 2.5 mm Cu) orthovoltage beam at 50-cm SSD. Sequential HVLs after the first are often larger since the lower energy, less penetrating portion of the orthovoltage x-ray spectrum, is preferentially filtered out. High-energy photons (e.g., 18 MV) scatter at smaller angles, but their Compton electrons have a large range, so their dose spread is greater than at 6 MV. The dc voltage applied to the x-ray tube of orthovoltage equipment is generated by a high-voltage transformer and rectifier circuit. Use of three-phase power, full-wave bridge rectification, and capacitor filtration are often employed to smooth and maximize the voltage applied to the x-ray tube. The center tap of the high-voltage transformer is often grounded,
which reduces electrical insulation requirements by allowing the x-ray tube anode to operate at +V/2 with respect to the ground and similarly, the cathode at - V/2 for an overall generating voltage V. Orthovoltage equipment is not in widespread use and will not be treated further. Additional characteristics of orthovoltage beams and equipment are described by Johns and C ~ n n i n g h a m . ~ ~
MEGAVOLTAGE X-RAY THERAPY Megavoltage x rays, as shown in Figure 2-1, typically deliver a 10-50 percent dose at the surface, reach a maximum at a few millimeters to several centimeters depth below the surface, and then attenuate less rapidly with depth than orthovoltage x ra~s.26~33 The skin-sparing effect of this low-surface dose coupled with slow attenuation versus depth facilitate delivery of high doses to deeply lying tumors. This is an important advantage of megavoltage x rays, since dose to overlying normal tissues (and especially the vascular layer of the skin at typically 1-5 mm depth) frequently limits the dose that may be delivered to a tumor. In addition, the greater depth dose permits directing several beams at the tumor from different directions, with dam-
TOTAGBODY AND HEMIBODY X-RAY THERAPY (MAGNA-FIELD THERAPY)
age to overlying tissues being further reduced. A depth dose of 67 percent for a 10 X 10 cm field at 10 cm depth with the dose maximum at 1.5 cm depth is representative of a 6-MV x-ray beam at 100-cm SSD. Similarly, 72 percent and 2.4 cm depth, respectively, are for 10-MV x rays. For 24-MV x rays at 100-cm SSD, an 83 percent depth dose at 10 cm depth, and a 4 cm depth of dose maximum is characteristic. The isodose distribution for a 6-MV x-ray field is shown in Figure 2-2. Figure 2-3 illustrates the four major component units of a representative megavoltage treatment unit, the Clinac 18. They consist of (a) gantry mounted accelerator and stand, a treatment couch, a modulator that powers the accelerator, and the control console with card rack housing associated printed circuit cards for electronic control functions. Table 2-1 is an abbreviated list of x-ray and electron beam performance specifications for the Clinac 18. The Clinac 18 is often referenced as much data and experience are available. Several references pertain specifically to the Clinac 18 treatment unit.5,7,18,43 Typically, x-ray energies of linac treatment units range from 4 to 25 MV with occasional higher energies to 50 MV. The lower x-ray energy (e.g., 4-8 MV) units, treating shallow lying tumors in the head, neck, extremities and other organs, are the workhorse units of radiotherapy departments. The higher x-ray energy (e.g., 10-25 MV) units are employed for deeper lying tumors in the body trunk such as the pelvis. Although more costly, multi-x-ray energy units provide both options, in one treatment unit (see Chap. 11). The optimal x-ray energy will also depend on the depth dose buildup and exit beam characteristics that influence the deposition pattern of dose in tumor and surrounding tissues. The flatness and symmetry of treatment beams is a measure of their departure from idealized rectangular contours, a view that simplifies our thinking and treatment planning processes. Wedges are beam modification accessories that tilt dose contours to compensate for a nonperpendicular entrant body surface. Similarly, tissue compensators are employed for more complex surface contours. Heavy, thick shielding (sometimes called shadow) blocks are often employed to shape a treatment field to the contour of an individual patient's tumor. Electron beam therapy is commonly employed for shallower tumors extending to the body surface and for boost doses. Compare the central axis depth dose curves for x-rays (Fig. 2-1) with Figure 2-4 for electrons, and also the isodose distributions for x-rays (Fig. 2-2) with Figure 2-5 for electrons. Often the square electron applicator fields are supplemented by a 1.2-cm thick shielding insert shaped for the individual patient. The production of treatment beams is described in more detail in Chap. 8 and treatment beam application accessories in Chap. 12. An extension of multiple field therapy, as noted above, is photon arc therapy. It provides a continuous change in beam direction by rotating the isocentric gantry over a preset angular range while directing the beam at the target volume. Most isocentric treatment units incorporate a 100-cm source axis distance (SAD). Dose rates at the isocenter may be varied and are typically 200-400 cGy/min at 1 m SAD, where the depth of overlying material isjust sufficient to placed,,,,, at the isocenter.
35
Megavoltage radiographs, called port films, taken with the patient in a treatment position, are used to establish correct patient positioning in relation to the field shape.l.7.64 The diagnostic quality of accelerator port films is a function of photon energy and anatomic detail visualization for megavoltage beams is poor, particularly at 10 MV and above. The visualization problem with high-energy beams is their high penetration, hence, small attenuation difference between bone and soft tissues, and consequently, very poor contrast ratio. The photon absorption process in megavoltage therapy is primarily due to Compton interactions. Their frequency of occurrence is closely proportional to the physical density of the absorber, and megavoltage port films primarily reveal density differences of the anatomy. Some high-energy units also provide a lower energy (4-6 MV) capability, for therapy and for obtaining better quality port films. A diagnostic voltage x-ray generator has been incorporated in a high-energy treatment unit.7364 The absorption process for the low-energy photons used in diagnostic radiology is primarily photoelectric. Photoelectric absorption is inconsequential even in calcium (bone) above about 0.2-MeV photon energy. This interaction is strongly dependent on atomic number and photon energy, and diagnostic voltage films have a richness of detail not found in megavoltage port films. The absorption of 80-kVp x rays per unit depth may be 100 times greater than for 6 MV in tissue.
TOTAL-BODY AND HEMIBODY X-RAY THERAPY (MAGNA-FIELD THERAPY) Very large megavoltage x-ray fields (magna-fields) are used to treat large portions of or the entire body volume. Various aspects of this technique have been reported by a number of investigators.'.l8-19.~2.30.31,~7.57.63 Such therapy is often identified as total-body irradiation (TBI), hemibody irradiation (HBI), partial-body irradiation (PBI), total-lymphoid irradiation (TLI), or total-nodal irradiation (TNI). The dose is severely limited for such large fields by normal lung tissue tolerance, with pneumonitis a potential clinical complication. High dose TBI is frequently used for immunosuppression in bone marrow transplantation as well as for treating lymphomas. It has also been employed for the treatment of rheumatoid arthritis and lupus. These magna-field techniques may constitute the primary treatment or be adjunctive for the latter diseases. Frequently, they are adjunctive for chemotherapy. The various methods used to irradiate TBI fields, which may be 2 m in length, have been described by Van Dyk63 and by Shank.57 A large treatment room permitting SSD values from 3 to 6 m may be employed. The patient is positioned in two or more orientations with respect to the treatment beam that may, itself, have several orientations. Dosimetric considerations for such therapy include dose uniformity, dose rate, and point of dose specification together with the fractionation
36
CHAPTER 2. RADIOTHERAPY MODALITIES
FIGURE 2-3 . Major coniponent units of a represfntative treatnlent unit, the Clinac 18, providing one x-ray energy of 10 MV and five electron energies ranging fronl 6 to 18 MeV (courtesy of Varian kqsnciates). (a) Gantry with stand and couch (b)modulator cabinet. (c) Control console with card rack.
MEGAVOLTAGE ELECTRON THERAPY
Depth In Water (cm)
FIGURE 2 4 . Central axis depth dose curves for 10 x 10-cm electron beams in water, 6-22 MeV in energy at 100-cm SAD.
regime. Dose rates range from about 5 to 50 cGy/min at the extended treatment distance. Total dose ranges from about 750 to more than 1000 cGy and lung shielding is frequently employed. The lower total dose values are delivered in one fraction prior to bone marrow transplantation, and treatment times as long as several hours may be required. Dose uniformity within + 10 percent is generally achieved, usually with the use of bolus and compensators. Linac beam energies used for this technique have primarily been 6 and 10 MV. A unique total body irradiation facility employing parallel opposed fields has been constructed using two 4-MV x-ray units.39
37
capability. Variable dose rates up,to 400 to 500 cGy1min are easily obtained, since the accelerator beam current required is typically at least several hundred times less than for similar megavoltage photon dose rates. This occurs because of the conversion inefficiency of producing x rays in the target, losses in the flattening filter, as well as the greater ionization of electrons occurring in tissue per unit depth when compared to x rays. At energies of about 25 MeV and above, the shape of the electron depth dose curve tends increasingly to resemble that of megavoltage photon beams. The therapeutic depth for electron beams is often defined as the depth of the 80 percent dose on the falling portion of the depth dose curve. In centimeters of water, this depth is very approximately equal to one-third of the electron energy in megaelectron volts. Figure 2-3 illustrates a radiotherapy unit, the Clinac 18, providing electron beams of 6,9, 12, 15, and 18 MeV. The electron beam performance specifications are given in Table 2-1. Figure 2-5 shows a typical isodose distribution in water for a 20-MeV electron beam. Compared to megavoltage photon beams, megavoltage electrons interact more rapidly and are scattered more widely by the intervening air, by treatment head components, and by the body tissues. Hence, to reduce dose outside the treatment volume, electron beam collimation is preferably carried out at or near the skin surface where the divergent effect of scatter is reduced. Although some continuously variable collimators are employed, fixed field applicators are customary. Electron applicators attach to the accessory mount and provide a range of discrete field sizes from about 5 X 5 cm to 25 X 25 cm. Irregular field-shaping inserts of a
MEGAVOLTAGE ELECTRON THERAPY Megavoltage electron therapy is used to treat tumors located near or extending to the skin surface and for boost therapy. Many articles have been written describing this technique. A number of these articles are cited herein.5926329-32333938,43761.65 Figure 2-4 illustrates central axis depth dose curves for electron beams for energies of 6 to 22 MeV. Compared to megavoltage x-ray beams the penetration is shallower and the surface dose is higher, typically 80-95 percent. The depth dose tends to FIGURE 2-5 . Isodose curves for a 20-MeV electron beam in water, 10 fall-off rapidly beyond the dose maximum, but more slowly as X 10 cm at 100 cm SSD. The applicator defining the field size is located 5 the energy is increased. Treatment energies range from about 4 cm from the water surface. Note the effect of electron scatter in widening the beam with increasing depth. to 25 MeV, although some accelerators provide 35-50 MeV
38
CHAPTER 2. RADIOTHERAPY MODALITIES
TABLE 2-1
. Clinac 18 performance specificationsfor x-ray and electron beams
X-ray beam
Electron beam
Energy
Energy
10 MV
6,9,12,15, and 18 MeV
Dose Rate
Dose Rate
100-500 cGyImin at 100 cm
100-500 cGy/rnin at 100 cm
Field Size
Field Size
OXOcmto35 X35cmatlOOcm continuously adjustable
4 X 4cmto25 X 25cmat100cm
Flatness
Flatness
5 3 % at 10 cm depth over 80% of both axes at 100 cm
25% at depth of Dmax. over 80% of both axes at 100 cm
Symmetry +2% at 10 cm depth of integrated doses to
Symmetry
longitudinal and transverse halves of the field at 100 cm Spot size
+2% at depth of dm,, to longitudinal and transverse halves of the field at 100 cm X-ray contamination , ; ;
8 .
FIGURE 3-8 . Schematic of a source with an internal impedance Rs connected to a msistive load RL
I(z = 0 ) = 1 (V+ - V-)ejm zo (3-14) We define the reflection coefficient p as the ratio of voltage in the reflected wave to that in the incident wave.
.- - - - - -
'. ; , ' ' ' ' .---. , , ;, i iiii - - _--i i, t; !\ ! i.---', ; ; _ _ - .' - , , .- ----- -" : : '-------_
1
a
I
I,
I
I
I
\
3
\
,
'
I
;----
'
I
,
4
, 8,
I,,
\
8
'.- - - _ _ _ _ _ _I-I-
I
- _ _ _ -,-'
-.
-b--
-b-.
,
1 I
FIGURE 3-7
----+----
, ', ; b , I
---___
\\
--+--
I-
,---*---.', \
e I ,
,
c I
\
, $
,
$ , 0
-*--
; 8 I ,
, I
I
. Summary of electromagneticfield distributionsfor circular waveguides (from Ref. 5).
55
RESONANCE AND RESONANT CAVITIES
FIGURE 3-9
Schematic of a low-loss transmission line.
z= 0 (3-15) FIGURE 3-10
since V(z = 0) = ZL X I(z = 0). One can also derive an equation for the input impedance at z = - 1 looking toward the load as follows: ZL + jZo tan pl
zi = zox Zo + jZL tan pl If ZL = Zo, no power will be reflected since V- = 0 and for this condition it is said that the transmission line is matched to the load. If ZL # ZO,then the voltage envelope along the transmission line will look as shown in Figure 3-10. We define the voltage standing-wave r a t i 4 ~ as~the~ ratio ~ )of the maximum voltage amplitude to the minimum voltage amplitude. VSWR
FIGURE 3-11
-
Vm, Vmin
= --
Iv+l+ Iv-l
1 + lp\ IV+ - IV-I - 1 -
I
Examples of mechanical resonance.
IPI
(3-17)
- Voltage waveform with ZL # ZOterminated with ZL.
RESONANCE AND RESONANT CAVITIES Resonance or resonant phenomena can be observed in everyday life. For example, in most musical instruments, such as the piano or violin, strings of discrete length vibrate at certain frequencies, and this phenomenon is called resonance. Sirnilarly, atoms or molecules can resonate at much higher frequencies. Figure 3-11 shows three mechanical resonant phenomena we may encounter, which can be visually observed in everyday life. In mechanical resonance, kinetic and potential energy are transferred back and forth. In electrical resonance, which we cannot visually observe without the help of sophisticated instruments,the electric and magnetic energy are interchanged at the resonant frequency f. Figure 3-12 shows simple lumped element resonant circuits for two different cases, that is, (a) series resonant and (b) parallel resonant circuits. One can
56
CHAPTER 3. MICROWAVE PRINCIPLES FOR LINACS
0
( a ) Series resonance
FIGURE 3-12
(b) Parallel resonance
. Lumped element circuits. (a) Series resonant and (b) Parallel resonant circuits.
derive second-order differential equations for these circuits in terms of the voltage and current.
The term "Q," which stands for "quality" or "quality factor," concerns relative energy damping and is often used in resonant circuit analysis. For a resonator with resonant angular frequency o , Q is defined as follows:
(3-18) . .
d2v Cp,+G dt
dV
-+-=O 'dl
Q=oX
V
\Lp
+ A2 sin w,t)
V = Voe-apt(B, cos opt + B2 sin opt)
(3-26)
(3-19)
The general solutions of these equations have the following form: I = Ioe-aSt (Al cos w,t
energy stored in the circuit energy dissipated in the circuit per second
The stored energy in the series resonant circuit is %LSP and the average power is m s P . Therefore,
(3-20) (3-21)
The stored energy in the parallel resonant circuit is lhCpVZ and the average power dissipated is l/(ZGpV2.Therefore,
where
= Lp = L; C, 1 R, = - = R and a d, the lowest resonant mode is TMo,,,. This mode can be considered as a TMol mode in a circular waveguide operating at cutoff. The electromagnetic fields are given as follows:
H
+ -
~ E o -J , (kr) 11
where k = 2.40510, q = intrinsic impedance of the medium and equals 120 x ohms for vacuum. The terms J, and J , are zero-
58
CHAPTER 3. MICROWAVE PRINCIPLES FOR LINACS
order and first-orderBessel functions of the first kind. As shown in Figure 3- 17, the unique features of a TMonomode are that the electric field is parallel to the axis and has maximumintensity at the axis of the cavity. Therefore, the TMo,Omode is suitable for accelerating paraxial particles. Figure 3-17 indicates the displacement current flow (in dashed line) and the real current flow (in solid line). For n > 1, the cavity is divided into n resonant LC circuits as shown in the equivalent circuits. The energy stored, U,in the cavity at aresonance TMol, mode is given by
20 x
15 x N -
E
0
-r".-
N
C
N
5-
lox
N
where E, is the dielectric constant of the enclosed medium and Q is the permittivity in free space in faradlmeter
5x
The power loss P, is calculated as follows: 7 ~ R, a
P, = FIGURE 3-15 . Mode chart for a cylindrical cavity: Resonant frequency versus square of ratio of diameter to length for TM ri?_dTE fundamental and higher order modes (from Ref. 6). ~-
(a) TMolomode
(b) TM,,
(e) TMo,,mode
(0
FIGURE 3-16
,mode
TM,,,mode
E;
q2
:
J (ka)[d + a]
e
where R, is the surface resistivity and equals 2.61 X 10-7 in ohms for copper at a frequency f. Therefore the Q of the TMolomode for copper is given by
(c) TM,
,,mode
(g) TM1,,mode
(4
TM,,,mode
(h) TM,,,mode
- Direction of electric and magnetic fields for TM fundamental and higher order modes for cylindrical cavity (from Ref. 9).
59
PERIODIC STRUCTURES AND COUPLING
dz o Therefore the velocity of the wave propagation - = -. dt S Since this is the velocity of the constant phase, it is called the phase velocity and is often denoted as up.
o
up =
Axis
P
(3-39)
In order to conceptually understand the group velocity, consider that two waves with the same amplitude but slightly different frequency are added as follows:
TM010
where
-5 Axis
FIGURE 3-17 . Radial variation of axial electric field E, (r) and equivalent circuits for (a) the TMolo mode and (b)the TMozo mode (from Ref. 4 of Chap. 4).
PHASE VELOCITY AND GROUP VELOCITY -
This means a wave with phase velocity up is modulated with a low-angular frequency A o as shown in Figure 3-18. The velocity of this modulation wave is called the group velocity us and is given as follows:
Group velocity is usually the same as the velocity of energy travel.
PERIODIC STRUCTURES AND COUPLING
Consider a wave propagating in the forward direction described by a voltage function of -
Therefore;
-
v+ei(ot - Pd
(3-36)
A periodic structure is a system consisting of multiple discrete elements spaced in a periodic manner. For example, the crystal structure of a monoatomic lattice, such as silicon, is a periodic
If the phase (wt - Pz) is constant @, as Modulation
=-+%
o t - PZ = bc
'Carrier Wave
then
+w (3-38)
FIGURE 3-18
. Illustration of group velocity (from Ref. 14).
60
CHAPTER 3. MICROWAVE PRINCIPLES FOR LINACS
K
M
K
M
K
M
K
M
K
(a)
FIGURE 3-19 impedences.
. Periodic Structures (a) Mechanical analogy with periodic springs and masses. (b)Microwave transmission line loaded with lumped
vp* = fA,
5
c
(4
.
FIGURE 3-20 Instantaneous electric field in (a) a cylindrical waveguide and (b)a disk loaded cylindrical waveguide.
61
PERIODIC STRUCTURES AND COUPLING
structure. These mechanical periodic structures can be represented by a series of mass M and spring K combinations, as shown in Figure 3-19a. A microwave transmission line loaded with lumped impedance, Z, spaced at periodic intervals, d, as shown in Figure 3-19b, is often used in microwave active and passive devices. There are two basic important properties for all periodic structures: (a) the phase velocity is less than the velocity of light and (b) they transmit only in the frequencies of the pass bands and not in the stopbands. Figure 3-20a shows a simple circular waveguide of radius a and the instantaneous electromagnetic field distributions of the TM,, mode. This structure cannot be employed for particle acceleration since the phase velocity up,given by
kind of structure is often called a slow wave structure. If a transmission line with the phase velocity up,given by
is loaded with lumped shunt capacitances Co at periodic intervals d, the new phase velocity will be given by
This means that the energy of propagating electromagnetic wave is locally stored in the shunt capacitances, hence slowing down the wave propagation. Slow wave structures are very often utilized in microwave passive and active devices. They hav very important roles in slowing the phase velocity to mat h the velocity of the electrons, especially in traveling wave amplifiers or in accelerators, where the electron beam must strongly interact with electromagnetic waves. The characteristics of passband and stopband can be understood by plotting propagation constant P versus wave num-
!
is greater than the velocilty of light c. By introducing a series of annular disks, as shown in Figure 3-20b, the phase velocity will be reduced. Thus, this
Cavity 1
Cavity 2
(a)
FIGURE 3-21 . Electric coupling of two cylindrical cavities (a) via axial iris and (h) via capacitance Cc in an equivalent circuit.
Cavity 1
Cavity 2
(a)
FIGURE 3 2 2 . Magnetic coupling of two cylindrical cavities (a) via slots in the disk between cavities and (b) via mutual inductance M in an equivalent circuit.
62
CHAPTER 3. MICROWAVE PRINCIPLES FOR LINACS
6.
ber k, where k is given by k = o.The plot is called the k p diagram or the Brillouin diagram. As shown in Figure 3-20b, a disk loaded waveguide can be considered a series of cylindrical cavities coupled through their apertures. Figure 3-21a shows two resonant cavities with TMolomodes coupled through an aperture. Since the aperture is located near the axis of the cavity where the electric field strength is maximum, these cavities are electrically coupled. Figure 3-21b shows an equivalent circuit of this type of coupling in which the electrical coupling aperture will be represented in a coupling capacitance C,. Similarly, a magnetic coupling can be achieved through a magnetic coupling iris as shown in Figure 3-22a. The equivalent circuit of this type of coupling is shown in Figure 3-226, where M is a mutual inductance. The relation between M and self-inductances L, and 4 is usually written as follows:
where k is called the coupling constant between L, and &. When two identical cavities (C, = C2, L1 = &, and R1 = R2) of resonant frequency fo are coupled, the total resonant frequencies are not fo anymore, as shown in Figure 3-23. For tight coupling (large k), two clearly separated resonancesfi and f2 appear. The relation between the coupling constant k and frequencies is given by
resonant mode or the waveguide mode. Figure 3-24 illustrates the differencebetween a single cavity and two coupled cavities. Figure 3-25 shows three identical pendulums coupled by springs. One can imagine three different pendulum motions. They are 1. All pendulums move in the same direction synchronously. 2. Each pendulum moves in an opposite direction. 3. Two end pendulums move in opposite directions while the center pendulum does not move at all. The mathematical analysis of these pendulum motions is rather complicated, but one can estimate the force acting on each pendulum. For case I, there is no force on the springs and for case 11, the force on the springs will be maximum, while for case 111, the force will be somewhere in between case I and case 11. Therefore, the velocity of motion of the pendulums for case I will be lowest since three masses move together, while case I1 will be highest. Thus, the frequency of pendulum motion will be highest for case 11, lowest for case I, and the frequency of motion for case I11 will be somewhere in be-
Single Resonator
Coupled Resonator
MODE AND DISPERSION When N resonators are coupled, N different resonator operational modes appear. These modes are different from the cavity Amplitude
Amplitude
Amplitude
4
4
fo
-
FIGURE 3-23 Frequency spectrum (relative field amplitude vs. frequency) for various degrees of coupling: (a) fork = 0, (no coupling), the system resonance stays at fo. (b)For relatively light coupling, (k small), two resonance peaks appear. (c) For tight coupling, (large k), two resonances are further apart.
.
FIGURE 3-24 Mechanical analogy, equivalent circuit, frequency spectrum, and cylindrical cavity electric field direction for single resonator and for two coupled resonators.
63
MODE AND DISPERSION
Case I
Case II
Stretched
Compressed
----
Case Ill
u
f
u
No Motion
FIGURE 3-25
. Three coupled pendulums, showing different modes of oscillation.
tween. The operational mode for case I is called zero mode, that is, zero phase shift between adjacent pendulum motions. Similarly, case 11mode is called the 7t mode (180" phase shift between adjacent pendulum) and case I11 mode is called the 1~12mode. When N resonators are coupled (see Figure 3-26 for an example of N = 7), N different operational modes exist and each mode is defined as follows:
where n = 0,1,2,3 ,..., N - 1. Figure 3-27 shows the amplitudes of axial electric fields for the allowed modes of seven coupled resonators. The seven operational modes and resonant frequencies can be plotted on a mode diagram as shown by the circles in Figure 3-28. This diagram is often called the dispersion relation of resonators, and the phenomenon is called dispersion; that is, the phases of motion of pendulums vary with frequency. In the disk loaded waveguide, as shown in Figure 3-26, the phase velocity of the wave varies with frequency.
64
CHAPTER 3. MICROWAVE PRINCIPLES FOR LINACS
Multicavitv O~erationalMode
Amplitude
j
A
-
7 resonances
pass band
: Freq.
FIGURE ,-26 . An example of seven coupled cavities and their corresponding seven resonance frequencies in the lowest pass band of frequencies.
SHUNT IMPEDANCE AND TRANSIT TIME
where Eo is the accelerating field and is given by
A simple circuit theory states that the power loss PL in a resistance R is given by
where V is the voltage drop across the resistor. Similarly, we define the shunt impedance r of an accelerating cavity assuming the accelerating field is time independent, as follows:
where Vm is the maximum energy gain by a particle going through the cavity and PL is the power lost in the cavity. If the cavity length is L, then the shunt impedance per unit length is given by
Figure 3-29a shows an ideal pill-box cavity for a particle accelerated from left to right by passing through the cavity. The accelerating electric field E varies with time t as E = Eocos ot. In the real case, the finite radius beam aperture creates radial components of the E field as shown in Figure 3-29b, thus the axial electric field E depends on z. Assuming E is independent of z, the velocity u of the particle is constant, and it passes through the center of the cavity at t = 0, the energy gain V of the particle is given by
Since z = ut
65
SHUNT IMPEDANCE AND TRANSIT TIME
Frequency
II
Shift I
I
I
I
I
I
:Per Cavity
FIGURE 3-28
Dispersion diagram for seven coupled cavities.
Thus the effective shunt impedance, Z,, in which the transit time effect is considered, can be defined as follows:
Zeff =zT'=-
v P,
xL
Section A-A
+ Current I Charge +,a
EField b
+---'A' LI HField
c
FIGURE 3-27 . Amplitudes of axial electric fields for the allowed modes of seven coupled cavities (from Ref. 4 of Chap. 4). Section A-A
(3-54)
where
T=
sin(oL/2u) 0L/2v
(3-55)
and is called the transit time factor. This factor is always less than unity. From eqs. (3-51), (3-52), and (3-54), Z will be given by
(3-56)
-
Current I Charge +,-
a
t--
+---'A'
EField
HField
b
c
b
(4 FIGURE 3-29 . Wall current flow I, end plate charges, E field and H field in cylindrical (pill-box) cavity: (a) without coupling apertures and (b) with axial coupling aperatures.
66
CHAPTER 3. MICROWAVE PRINCIPLES FOR LINACS
If w e assume a n accelerator consists of a series of pill-box cavities each of length L = M2 and the particle velocity v = c, K K
the transit time factor will be only 0.637 ( = sin -/-),
2 2
and the
Z,, will b e only 0.406 Z. For a pill-box cavity of radius a and length L, the shunt impedance per unit length Z is given b y
z=
rl; L na R, ( L + a) J : (ka)
where q, = 1207r and R, = 2.61 X 10-7 X $(for copper). Since ka = 2.405 for a TMOlomode resonance, cavity radius a should b e 3.825 c m in order to resonate at the f r e ~ e n c of y 2998 MHz. If L = 5 c m (A = 10 cm), the shunt impedance per unit length Z will b e 146 M W m . Therefore the effective shunt impedance is 146 X 0.406 = 59 M W m . T h e transit time factor will increase if L is decreased, but Z will decrease. Figure 3-30 shows the relationship of Z,, and L. I n general, the accelerating cavities have beam holes, as shown in Figure 3-29b, and the accelerating electric field, E along the beam axis, is not uniform like the pill-box in Figure
3-29a. If E = Eo(z) cos wt, then the transit time factor should b e given by
. Effective shunt impedance Z? versus length L of cylindrical (pill-box)cavity.
FIGURE 3-30
0 2n v where y = - = -and p = -. v Qh c T h e most useful references for Chap. 3 are textbooks, which are listed in the reference section (1-21).
REFERENCES 1. Panofsky W, M. Phillips: Classical electricity and magnetism, New York, Addison-Weslley Publishing Co., 1962. 2. Purcell, EM: Electricity and magnetism. Berkeley Physics Course; New York, McGraw-Hill Book Co., 1965, vol2. 3. Slater JC, NH Frank: Electromagnetism. New York, McGrawHill Book Co., 1947. 4. Cheng, D.: Field and wave electromagnetics, New York, Addison-Wesley Publishing Co., 1983. 5. Ramo S, John R. Whinnery, T. Van Duzer: Fields and waves in communication electronics. New York, 1965. 6. Collin RE: Foundations for microwave engineering, New York, McGraw-Hill Book Co., 1966. 7. Feynman R: Lectures on physics: New York, Addison-Wesley Publishing Co., 1964. 8. Slater JC: Microwaved electronics. Van Nostrand Co., 1950. 9. Ishii TK: Microwave engineering, The Ronald Press Co., 1966. 10. Ginzton EL: Microwave measurement. New York, McGraw-Hill Book Co., 1957. 11. Laverghetta TS: Microwave measurements and techniques. Artech House Inc., 1975. 12. Slater JC: Microwave transmission. New York, McGraw-Hill Book Co., 1942. 13. Plonsey R, R Collin: Principles and applications of electromagnetic fields. New York. McGraw-Hill Book Co., 1961. 14. Brillouin L: Wave propagation and group velocity. New York, Academic Press, 1960. 15. Crawford, Jr. Frank S: Waves. Berkeley Physics Course, New York, McGraw-Hill Book Co., 1968, vol3. 16. Atwater, HA: Introduction to Microwave Theory. New York, McGraw-Hill Book Co., 1962. 17. Cardiol, FE: Intorduction to Microwaves. Artech House, Inc., 1984. 18. Stratton, JA: Electromagnetic Theory. New York, McGraw-Hill Book Co., 1941. 19. Moreno, T: Microwave Transmission Design Data. New York, McGraw-Hill Book Co., 1948. 20. Purcell,EM: Electricity and magnetism. Berkeley physics course; New York, McGraw-Hill Book Co., 1965, vol. 2. 21. Marcuvitz, N: Waveguide Handbook. New York, McGraw-Hill Book Co., 1951.
C H A P T E R
Microwave Accelerator Structures
In this chapter, the structures of typical microwave electron accelerators are described. Electrons are generated in an electron gun and injected into the accelerator. There are basically two distinctly different types of accelerators, namely, the "traveling-wave accelerator" and the "standing-wave accelerator". These two accelerators have advantages and disadvantagesand this chapter gives a detailed comparison of both types.
at ground potential. The current in the beam for a given anodeto-cathode voltage is largely determined by the ratio of the cathode-anode spacing and the cathode diameter. In order to vary the beam current over a wide range without varying the anode-to-cathode voltage, a control grid is incorporated between the cathode and the anode as shown in Figure 4-2 forming a triode, a three-electrode electron gun. The required grid voltage to control the beam current is typically 2-5% of the anode-to-cathode voltage.
ELECTRON GUNS AND INJECTION CATHODE Electrons are injected into an accelerator structure from an electron gun. A cross-sectional view of a two-electrode diode electron gun is shown in Figure 4-1. It consists of a spherically shaped cathode button 1,focus electrode 2, anode 3, and heater 4. The cathode is at a negative potential with respect to the anode. Electrons emitted from the cathode are accelerated and focused through the beam hole in the anode, which is usually
There are many types of cathodes in use for various gun applications. They are basically divided into two different types-namely, oxide cathodes and dispenser cathodes. The oxide cathode can be operated at relatively low temperaturebut the maximum current density is about 1 Alcmz. The advantages of dispenser cathodes are generally, higher current density at the cathode (100 AIcm2) and less susceptibility to gas poisoning. But the operating temperature for dispenser cathodes is much higher (1 100°C) than for oxide cathodes (800°C). Tung-
Heater @-
Focus ( Electrode Anode
FIGURE 4-1
. Cross-sectional view of a diode electron gun.
FIGURE 4-2 trol grid.
- Cross-sectional view of a triode electron gun with con-
4
68
CHAPTER 4. MICROWAVE ACCELERATOR STRUCTURES
sten dispenser cathodes are used almost exclusively in microwave devices, since they are capable of high current densities with long lifetimes. Oxide cathodes are used in low current applications, such as in cathode ray tubes. A tungsten dispenser cathode consists of a porous tungsten matrix impregnated with barium oxide (BaO), calcium oxide (CaO), and alumina (A1203)compounds with a proper molar ratio. Because of the high temperatures required, the surface coverage of barium decreases with time, which means, the cathode life is limited. A significant improvement in cathode life can be made by overcoating the tungsten surface with osmium, iridium, or ruthenium, thus reducing the work function by about 20 percent. Consequently,the cathode can operate at 100°C lower temperature, which increases the life of the cathode by a factor of about 10.
Equation 4-1 can be rewritten as I = P X V, where P is a geometrical function called "perveance," I is in amperes (A) and V is in volts, (V). Typical values of perveance for electron guns used in medical linear accelerators range from 0.1 x 10-6 to 0.5 X 10-6. Modern electron guns are designed using a combination of digital computer techniques and experiments using a beam analyzer. Figure 4 4 shows typical electron trajectories in a diode gun of perveance 1.4 X 10-6 (1.4 rnicroperveance unit) at 40-kV anode voltage.
ELECTRON INTERACTION WITH MICROWAVE FIELD MOTION OF ELECTRONS
DESIGN OF AN ELECTRON GUN Pierce type convergent electron guns are commonly used in linear accelerators and klystrons. The conceptual model of a Pierce gun uses an annular segment of the electron flow between two concentric spheres. The outer sphere comprises the cathode, and the inner sphere the anode. The formation of the electron beam using this concept is shown in Figure 4-3. The current from such a gun will be given by solving Poisson's equation. I = 29.3 X
1 - cos 8
a2
v"
where
Electrons are accelerated in the direction of motion only by an electric field E, while magnetic fields exert forces at right angle to the direction of electron motion and to the direction of the magnetic field B. This relation which is called the Lorentz force equation, is given by
where F is the force acting on an electron of charge e and velocity v. This equation says that an electric charge is acted on by two types of force: an electric force, independent of its velocity, and a magnetic force, proportional to its velocity u. Also, it says that the electric force is along the electric field line and the magnetic force is at right angles to its velocity and the magnetic field. In the MKS system of units the force is given in newtons, the charge in coulombs, electric field E in volts per meter, u in meters per second, and magnetic field B in webers per square meter (1 weberIm2 = 104 gauss). Since dP d F = - = - (mu), eq. (4-2) can be written as dt dt
V = voltage between cathode and anode 0 = cathode radius angle
d dt
- (mu) = e (E
+ v X B) Equipotential Lines
. Electron Beam
49
Cathode
Electrode
-z 40 kV Anode
\
Electron
Catl
,
.- - - - _ - -.,
FIGURE 4-3 . Pierce gun electron beam trajectories.
FIGURE 4-4 . Conlputer simulated electron trajectories in a diode gun. (Microperveance = 1.4. Voltage = 40 kV.)
69
ELECTRON INTERACTION WITH MICROWAVE FIELD
For relativistic motion,
where mo is the rest mass of an electron and c is the velocity of light. It can be shown that the kinetic energy T of an electron is given by T = (m - mo) c2. Thus Energy. Mev
FIGURE 4-5 . Relative velocity for an electron and a protron as a function of kinetic energy.
where: 1
P = -vc and y = (1 - P2)-2
(4-6)
For p 55 Transit time factor 0.82 Accelerating gradient (MVIm) 4.3 Normalized peak surface field 2.7 Group velocity (%) .012 Beam Hole Diameter (mm) 22 Fill time (ps) .37 Magnetron power (MW) at accelerator input 1.8
SW accelerator
CL-61100 6 rl2 SW
0.3 100 0.91 20 6.5 0.03 5 .35 2.3
86
TABLE 4-3
CHAPTER 4. MICROWAVE ACCELERATOR STRUCTURES
. General comparison of TW and SW accelerators ~p
TW accelerator
Shunt impedance Isolator or circulator Maximum accelerating beam current Tuning sensitivity Input coupler design Buncher design Spectrum sensitivity on accelerating field Coupler Coupler position
SW accelerator
Low Not needed
High Needed
High -2 A High Complex Rather complex
Low -0.5 A Low Simple Simple
Low Dual First and last
High Single Any
circulator, since it is a matched device. The maximum accelerating current can be as high as 2 A for a 2 m length TW accelerator. Also, the bunching characteristics of the TW structure are inherently less sensitive to the variation of the accelerating field. This means the T W accelerator can offer a relatively wide range of energy variation without sacrificing the quality of the beam spectrum.
DESIGN OF ACCELERATOR CAVITIES There are many important parameters of an accelerator cavity to consider in designing a highly efficient linear accelerator. They are frequency, shunt impedance, (Q), transit time factor, power loss, stored energy, and peak surface electric and magnetic fields. These values can be obtained by solving Maxwell's TABLE 4 4
TABLE 4-5
-
LALA computational examples
--
Axial average accelerating field Frequency Transit time factor Q Factor ZPlL Maximum surface field Stored energy Power loss
1 MVlm f = 3002.16 MHz
0.788 18,529 102.80 MWm 3.03 MVlm 0.77 X 10- J 485.8 W
equations with given boundary conditions. Since most of the accelerator cavities do not support pure TM or TE modes, it is almost impossible to use analytical techniques to solve them. Within the last two decades, computer programs have been developed and used to optimize cavity structures. Table 4-4 summarizes these programs in the order of the year developed. Recently, the 3-D computer codes such as MAFIA, have been extensively used in computation of nonaxisymmetric cavity structures. Figure 4-32 shows a typical S W accelerator cavity cross section, and Table 4-5 summarizes a computational example of a cavity used in the Varian Clinac 6X accelerator guide. As shown in Table 4-5, power loss, stored energy, and surface electromagnetic field measurements are given in terms of the normalized axial accelerating gradient of 1 MeVIm. One of the most important parameters for cavity geometry optimization is the effective shunt impedance per unit length, that is,
=. L
This quantity varies with beam hole radius R,, nose outer
radius R,, nose cone radii R R l and RR2, web thickness L,, nose height L,, and cavity radius R, (see Figure 4-32). Since
. List of programs for computing accelerator cavities Year developed
Developed
LALA
1965
H. Holt (Los Alamos)
Axisymmetric lowest mode (TMol) Finite difference method Rectangular mesh (constant mesh size)
SUPERFISH
1976
K. Halbach (LBL)
Axisymmetric higher mode (TMon)Triangular mesh (variable mesh density)
LACC
1978
A. Konrad
Axisymmetric higher mode (TMon) Finite element method Triangular mesh with variable mesh size
ULTRAFISH
1981
R.Gluckstern
Nonaxisymmetric higher mode (TMmn) Triangular variable mesh
URMEL
1982
T. Wciland (DESY)
Nonaxisymmetric higher mode (TM,) Rectangular mesh
MAFIA
1984
T. Wciland
Three dimensional finite difference code with PIC (particle in cell) modules allowing self consistent particle-field interactions
Name of program
by
Features
DESIGN OF ACCELERATOR CAVITIES
the resonant frequency f should be kept constant during optimization of the cavity geometry, these dimensional quantities cannot be varied independently from each other. For instance, if R, is increased (increase inductance), then L,, must be decreased (decrease capacitance). Figure 4-33 shows the effective shunt impedance as a function of beam hole radius. A 1-mm reduction of beam hole radius results in 272 improvement of greater than 3 percent. ZT2 Figures 4-34 and 4-35 shows -as a function of web thickL ness L, and nose height L,. In all computations, the cavity radius R, was adjusted to keep the resonant frequency at 3 GHz. In practice, the actual observed shunt impedance is about 15percent lower than the computed result.. This occurs because we do not include the effects of surface roughness, coupling slots and iris, coupling cavity power loss, finite beam energy spectrum, and wake field excitation by the beam.
87
REFERENCES 1. Lapostalle PM: Linear accelerators. North Holland Publishing Co., 1970. 2. Septier A: Focusing of charged particles. New York, Academic Press, 1967, vol 1 and 2. 3. Livingston M, Blewett J: Particle accelerators. New York, McGraw-Hill Book Co., 1962. 4. Humphries Jr, S: Principles of charged particle accelerators. New York, Wiley, 1986. 5. Livingston MS: Particle accelerators: A brief history. Harvard University Press, 1969. 6. Chodorow M and Susskind C: Fundamentals of microwave electronics. New York, McGraw-Hill Book Co., 1964.
C H A P T E R
- 5
Microwave Power Sources and Systems
The various microwave components that are used to generate the multimegawatt pulses of rf power and to transport them to the accelerator guide are discussed in this chapter. Medical electron accelerators are usually operated in the S-band at 2998 MHz (10cm wavelength) or 2856 MHz (10.5-cm wavelength). The dimensions of the microwave components are of this order, that is, in the general region of 10 cm. The capability of a microwave source to produce peak power varies approximately as the square of the wavelength. Thus, operation at an X-band (-3-cm wavelength) to permit shorter accelerator structures has been limited by the availability of reliable multimegawatt rf sources. However, this is in the process of change, since such sources are becoming available for physics accelerators. Operation at the L-band (-23-cm wavelength) has not been practical or desirable for medical accelerators because the accelerator structure and the microwave components become too large and there is no need for the high rf power capability of L-band. The rf sources for medical accelerators are usually magnetrons for low and medium energy machines and klystrons for high energy machines. The high peak rf pulse power needed for high energy machines is more readily and reliably obtained with a klystron than a magnetron. This occurs because the functions of electron emission, cavity-beam interaction and spent electron beam power dissipation are separated and distributed in the klystron but are all wound up in one small volume in a magnetron. The magnetron is a self-oscillator. It is kept tuned to the natural frequency of the accelerator structure by feedback (automatic frequency control, AFC) of an rf signal from the accelerator structure electromagnetic field to a motorized plunger in the magnetron cavity array. The klystron is used as an amplifier. Its frequency is determined by that of the rfdriver, which is usually an essentially all solid state oscillator amplifier, perhaps with a cavity stabilized microwave grid-tube output stage. The rf source needs to be isolated from rf power reflections from the accelerator structure. This need is greater with SW than with TW structures. Fenite microwave devices (circulators) are used to divert such reflected power away from the rf source into a water cooled rf load. A rectangular waveguide, pressurized to hold off the high rf electric fields, is used to
transport the rf power. Radio frequency ceramic windows in a short cylindrical waveguide section are used to separate the pressurized section (power source side) from the vacuum (accelerator and rf source). A rotary joint in a short cylindrical waveguide section is used in the connection between the stationary stand and the rotary gantry. Various other waveguide components are used to change direction or divide the flow of rf power. The cross section width of rectangular waveguides and related components is somewhat larger than one-half wavelength, in order to transmit the fundamental frequency rf power efficiently. These various types of microwave sources and components are described with greater clarity in the following subsections.
MAGNETRONS Magnetrons are frequently used as a source of microwave power for linacs, particularly for lower energy x-ray treatment units. The diode type magnetron was invented in 1913 by Arthur Hull. It is confined to short wave frequencies and requires an external LC resonant circuit. The cavity type magnetron was invented by Boot and Randall in 1940. Both the output power and frequency were markedly increased compared to previous split anode type magnetrons, making high definition radar possible during World War 11. Such a device is placed in a static magnetic field B, which is perpendicular to the plane of Figure 5-1. The electrons emitted from the inner cylindrical cathode are drawn toward the concentric anode by the positive anode potential. Surrounding the cathode is an array of small cavities linked together so as to form a slow wave structure. Figure 5-1 shows the cross section of electron trajectories, the dc pulsed electric field Edc,the magnetic field H, and the wall current I in the individual cavities of a magnetron. As soon as the electrons acquire a velocity, they are subjected to the Lorentz force e(u X B), and are turned in a tangential direction, as shown in Figure 5-la. The individual electrons perform a complex cycloidal motion around the cathode. Ro-
90
CHAPTER 5. MICROWAVE POWER SOURCES AND SYSTEMS
tial build up of oscillation is obscure, experimental results suggests that the dense space-charge cloud becomes unstable and causes preoscillation before full oscillation. In order to have a smooth transition between these states, the rise time of the anode dc bias should be at a limited rate (100 kV/ps). The performance of a magnetron in practical operation can be characterizedby two different diagrams. They are known as the performance chart (see Fig. 5-2) and Rieke diagram (see Fig. 5-3). Several parameters determine the operational performance of the magnetron. They are (a) the magnetic field B; (b) the anode current I,; (c) the anode voltage V,; (d) the load impedance Z,; (e) the output power Po; and (f) the operation frequency f. Figure 5-2 shows the performance chart of the English Electric Valve (EEV) magnetron 5125. This chart shows curves of constant magnetic field, peak output power, and efficiency versus peak cathode voltage and current. The dependence of overall efficiency on current for a constant magnetic field is related to the extent to which the rotating space charge spokes depart from synchronism with the phase velocity of the rf field on the anode structure. This chart does not show that the operation frequency varies with the anode current. This phenomenon is known as frequency-pushing and the value of AJAI, atf = fo, the resonant frequency, is called the "pushing figure". The Rieke diagram of Figure 5-3 shows the dependence of Magnetic Field (Grauss)
- - - Peak Power Out (MW) - . - . - . - Efficiency (%)
FIGURE 5-1 . Cross section of cavity type magnetron, showing (a) electron trajectories, wall current I, rf magnetic field H, and dc pulse electric field: Ede. (b)Radio frequency electric field Er/and rotating electron space charge cloud. (From Ref. 10)
tating spokes of electrons, as shown in Figure 5-lb, are formed due to the influence of space charge forces, and interaction with the rf cavities. This action induces the oscillatingrfelectric fields Edin the resonant cavities. the^ mode, in which the phase differencebetween adjacent cavities is .rr radians, is most commonly used. When the rotating spokes of electrons are traveling at the same velocity as the phase velocity of the mode, a strong interaction between the rf electric field at the entrance of cavities and electrons takes place. As the rf fields act so as to remove energy from the moving electrons,the reduction in kinetic energy of the electrons will be converted into the high-frequency electromagnetic energy. The electron will take up a new orbit of smaller radius and will spin back into the cathode, heating it, a phenomenon called back bombardment. Though the exact mechanism of ini-
Peak Cathode Current (A)
FIGURE 5-2
. Performance chart for EEV 5125 magnetron.
91
KLYSTRONS
minimum ratings of operating conditions for the 2.6 MW EEV 5193 magnetron. These magnetrons are mounted in a gantry such that the tuner axis is parallel to the axis of gantry rotation to minimize frequency variation from this source. The heater power is sewoed to a lower value as the rf power is increased in order to protect the cathode from over heating by back bombardment from electrons as described earlier.
KLYSTRONS
FIGURE 5-3
. Rieke diagram for EEV 5167 magnetron. (Pulse width;
4ps PRF; 250, Magnetic Field; 1550 G, Anode curret; 110 A.)
output power and frequency on the load impedance. The polar coordinate system consists of concentric circles, which give the reflection coefficient in terms of the VSWR [see eqs. (3-15) and (3-17)] and the straight lines give the phase position of the first minimum of the electric field from a reference plane. Table 5-1 lists the performance specifications of three EEV magnetrons used for Varian Clinacs. Table 5-2 shows the maximum and TABLE 5-1
. English electric valve magnetron specifications
Machine model
Typical peak input voltage vp OtV) Peak input current 1, (A) Average input power Pi (kW) Peak output power Po(MW) End of life power Po end (MW) Maximum duty Maximum PPS
CL-4 CL-6X L-200 43 (44 max)
CL-61100
47 (48 max)
CL-41100 L- 1000 L-3000 47 (48 max)
90 (100 max)
105 (110 max)
105 (110 max)
5.o
4.4 (4.7 max)
7.O
2.0
2.6
2.6
1.7
0.00116 323
2.4
0.00085 230
2.4
0.00135 323
The klystron was invented in 1937 by the Varian brothers. The klystron became a precursor for the development of the multicavity magnetron in England. It was ironic that the early klystron did not exhibit high power and efficiency comparable to the magnetron, which was employed in the World War I1 radar transmitter. Therefore, the klystron was destined only for local oscillator use in wartime radarreceivers. Since World War 11, the klystron amplifier has undergone a spectacular evolution, and it has become one of the most widely used devices for the amplification of microwave signals and can provide multimegawatt power output. Figure 5-4 shows a cross-sectional representation of a typical klystron amplifier. It consists of three major sectionsthe electron gun, four rf cavities, and the collector section. An electron beam, formed at a high potential, leaves the electron gun and passes into the first cavity (buncher cavity), which is excited at approximately its resonant frequency by the input rf signal. In each rf period, late arriving electrons are accelerated and early arriving electrons are decelerated by the axial electric field of this cavity. This causes the late electrons to catch up with the early electrons in each rf period as they drift towards the output rf cavity. The effect of this velocity modulation on electrons of the beam is to produce a bunching of the electrons into clusters along the beam axis. These bunches of electrons will induce rf current within the output cavity (catcher cavity), producing the TABLE 5-2
EEV 5193 magnetron operating conditions
Magnetic field Heater voltage Heater current (surge) VSWR (load) Pulse width Rate of rise Outlet water Temperature Water flow Anode voltage Anode current
Minimum
Maximum
Unit
1000 8.0
1575 10.0
G V
20 1.5 4.5 120
A
50
"C GPM kV A
1.06 42 60
48 110
PS kV/ps
92
CHAPTER 5. MICROWAVE POWER SOURCES AND SYSTEMS
rf
rf Output
Buncher Tube
1
Buncher
Heater Cathode
+- Electron Gun FIGURE 5-4
I I I
First Cavity 1st 2nd (Buncher) Intermediate Cavities
1
Catcher
Collector Last Cavity (Catcher) I I I
rf Section I
Collector+ I
. Cross section of a typical klystron amplifier.
rfpoweroutput from the tube. Figure5-5 illustrateshow abunch of electrons,passing through the output cavity, generates an oscillating rf current on the cavity wall. As the bunch of electrons approaches grid 1, the free electrons in grid 1 will be repelled, since negative charges repel each other, as shown in Figure 5-5. These repelled electrons flow from grid 1 toward grid 2, generating an rf displacement current across the gap from grid 2 to grid 1. Similarly,as the bunch of electrons approaches grid 2, a reverse rf current is generated. When the sequential bunches of electrons traverse the output cavity gap, with a time interval between them equal to the time interval of one cycle of the rfresonant frequency of the cavity, a strong interaction will take place and lead to the generation ofrfpower. Since theelectron beamis delivering energy to the cavity (conversion of kinetic energy to electromagnetic energy), it is slowed in velocity. The beam thereforeamves at the collector with less total energy than it had when it passed through the input cavity. This difference in electron beam energy is approximately equal to the rf energy delivered from the output of the cavity. The residual kinetic energy of the beam is dissipated as heat in the collector. Figure 5-6 is a schematic cross section of a high power, four-cavity klystron. The effect of the two intermediatecavities is to improve the bunching process. The partially bunched beam excites the intermediate cavity and is further bunched by the rf field. The result of the additional cavities is to increase amplifiergain, efficiency, and power output. The electron beam can be focused by the axial magnetic field of the magnetic focus coils, as shown in Figure 5-6. Figure 5-7 shows a cutaway view of a high power four-cavity pulsed klystron amplifier similar to the one used in the Varian Clinac 1800. Table 5-3 shows the typical operating condition of this klystron. The operating frequency can be varied by adjusting the diaphragms in the buncher and intermediate cavities, which are shown in the enlarged view in Figure 5-7. Radio frequency drive is introduced from a coaxial line by the coupling loop in the input cavity.
-
FIGURE 5-5 Excitation of rf wall curmnts in klystron intermediate or output cavity by passage of electron bunch.
Table 5-4 summarizes the comparison of S-band magnetron and klystron tubes. The peak output power of the klystron can be extremely high (recently, a 60-MW klystron develop ment was completed at SLAC), but it is bulky and requires very high voltage to drive. Thus, it cannot be placed in the rotational gantry.
RADIO FREQUENCY DRIVERS The rf driver is a stable, tunable rf source used to drive a high power klystron amplifier. For medical linacs, a pulsed rf driver is used with peak power around 300 Wand nominal pulse width of 12 ps. The rf frequency can be locally set and remotely varied about the set point. The peak power can be remotely
93
RADIO FREQUENCY DRIVERS
TABLE 5 4
Output Window
Function Output power Life Cost Magnet Operation Voltage Insulation
Electron Bunch Output Iris Output Cavity (catcher) Third Cavity Second Cavity lnput Cavity (buncher)
4 ! Electron Beam
L::!::\* .. ....... .. .
FIGURE 5-6 stron.
TABLE 5-3
Oscillator Low (3 MW) Short (2000 h) Low Permanent Moving Low (45 kV) Potting
Amplifier High (7 MW) Long(10,OOO h) High Electromagnet Stationary High (140 kV) Oil tank
programmable. The pulse repetition rate (prr) can be controlled by an externally supplied trigger pulse, and the pulse width can be adjustable. Table 5-5 lists the specifications of the 2856MHz rf driver used for the CL-1800. Frequency stability is one of the most important criteria for stable operation of the linac. Compared to the magnetron system, improved frequency stability arises in the klystron system where the frequency determining and amplification functions are separated. TABLE 5-5
. Radio frequency driver specifications
1. Frequency Center frequency Local tuning range Local tuning sensitivity
A"ode
Heater
2. Output power Peak power Isolation Adjustment range 3. Amplitude stability Short term (droop and jitter) Long term (flat top)
Clinac 1800, (VA8252) klystron operating
parameters --
Mode
Klystron
Min.
- Schematic cross section of a high power four-cavity kly-
-
Magnetron
lnput Loop
. . . :.', ;.:.:.:.:
Cathode
. Comparison of S Band magnetron and klystron
Low mode
+
Frequency 2856 2.5 Peak output power 3.0 Peak beam power 7.5 Gain at saturation 47 1.2 Load VSWR Beam pulse width 5.8 Repetition rate (max) 360 Peak beam voltage 110 Peak beam current 72 Heater voltage 7.5 Heater current 30 3 Warm-up time 10 Efficiency 43
+
High mode
Unit
+
MHz MW MW dB
2856 2.5 5.5 11.5 50 1.2 5.8 180 127 92 7.5 30 + 3 10 53
PS Hz kV A Vr.m.s. Ar.m.s. min %
4. Frequency stability Deviation and pulling Transient deviation rate Warm-up drift Frequencyltemperature coefficient 5. Pulse rate P'= Width (start to stop) Turn on time 6. Others VSWR of load Input voltage Trigger in pulse amplitude Trigger in pulse duration
Nom.
Max.
Unit
MHz MHz MHzItum W dB %
% %
kHz kHds kHz kHzf'C PPS FS PS
vrms
v PS
FIGURE 5-7 , (a) Cut-away four-cavity klystron, similar to that employed in the Clinac 18. Views (b)and (c)are cut away individual cavity sections. (b)Enlarged view of the bottom cavity, the input power coupling loop is on the right and a fine tuning device is on the left. (c) Enlarged view of cavity number three; the fine tuning device has been cut away in this view.
CIRCULATORS
FIGURE 5-8
95
. Block diagram of klystron rf driver employing triode output stage.
Figure 5-8 shows the block diagram of an rf driver. It consists of a voltage controlled oscillator (VCO) and frequency multiplier, S-band phase locked oscillator (PLO), positive-intrinsic-negative (PIN) diode modulator, S-band amplifier and final triode amplifier. The VCO must have a high-frequency temperature stability of 10-5 MHzPC. The S-band PLO has a phase lock range of 2 6 MHz and provides a low noise, stable rf signal. The PIN modulator modulates the rf signal by the modulation pulse (360 pps, 1 2 ~ ssupplied ) from the modulator drive circuit. This is amplified by a three-stage transistor Sband amplifier to 30 dBm (1 W). This pulse modulated signal is amplified by a triode amplifier to between 55 (300 W) and 57 dBm (500 W). An isolator is provided at the final output end to prevent damage due to external load variations.
CIRCULATORS A circulator protects the rf source from rf power reflected back to the source. Circulators are either three or four port devices using ferrites, a magnetic material which rotates electromagnetic fields. A three port circulator has the property that a wave incident in port 1 is coupled into port 2 only; a wave in port 2 is coupled into port 3 only; and so on. The ideal circulator is an impedance matched device; that is, with all ports except one
terminated in matched loads. The input impedance of the remaining port is equal to the matched load. Figure 5-9 shows the schematic diagram and cross section of a Y junction waveguide type three-port circulator. The ferrite disks are magnetized by a static magnetic field B, which is applied perpendicular to the figure shown and gives the junction the required nonreciprocal property. This can be understood as follows. A Y junction circulator behaves like alow loss transmission cavity. The resonance of the circular disk is in the dipolar mode in which the electric field vectors are perpendicular to the plane of the disk (parallel to B) and the rf magnetic field vector lies parallel to the plane of the disk. Figure 5-9a illustrates this dipolar mode of a ferrite disk in the unmagnetized case. The rf input power at port 1 will then divide equally between port 2 and port 3. When the disk is magnetized as shown in Figure 5-913, the mode defined by the E and H field pattern rotates 30" and port 1 will couple to port 2, but port 3 will be isolated from port 1. Figure 5-10 shows the TM,,, like mode pattern for the magnetized case where a wave incident in port 1 is coupled to port 2 only. Alternatively, any power entering at port 2 will set up a new mode that will result in this power leaving at port 3, and port 1 will be isolated, and so on. Typical values of insertion loss and isolation for waveguide type S band three-port circulators are 0.15 and 20 dB, respectively. A four-port circulator may be constructed from two magic Ts or hybrid junctions and a gyrator as shown in Figure 5-lla.
%
CHAPTER 5. MICROWAVE POWER SOURCES AND SYSTEMS
Port 3
Isolated Port
1
Port 3
-
FIGURE 5-10 Mode pattern of an H plane three-port circulator using TMiio mode.
Y
Output
FIGURE 5-9 . Dipolar modes of a ferrite disk and Yjunction threeport circulator in (a) an unmagnetized state and (b) a magnetized state. (From Ref. 4) The static magnetic field B perpendicular to the page is not shown.
The gyrator produces a phase shift of 180" for propagation in one direction but not for the other direction. Consider a wave incident in port 1. This wave is split into two waves, which are in-phase and equal in amplitude in the side waveguides b and d, with no power delivered to port 3. These waves are combined and delivered to port 2, since the gyrator does not affect the phase of the wave leaving waveguide b. Awave incident in port 2 will be similarly split into two waves, but this time with one arriving at d with a phase c$ and other arriving at b with a phase c$ + IT, because of the presence of the gyrator. The two waves cancel at port 1 but combine and emerge from port 3 in the hybrid junction (magic T). Figure 5-llb shows two different types of four-port circulator, where one of the circulators uses
--t
Gyrator
-
FIGURE 5-11 Afour-port circulator. (a) Conceptual view of a four-port circulator. (b) Two four-point circulators; one uses a 180' gyrator and the other uses two 90' gyrators. (From Ref. 6 and Ref. 17 of Chap. 3)
97
OTHER MICROWAVE COMPONENTS
To Accelerator
u Shunt Tee
I I
pO?Pofi Circulator H-Plane Bend
lan
RF Source
.
FIGURE 5-12 Schematic of a four-port circulator between the source and accelerator structure.
a 180" gyrator and the other one uses two 90' gyrators. The latter offers a more compact structure, but requires multiples of magnets forboth gyrators. Figure 5-12 shows schematically the use of a four-port circulator between the rf source and the accelerator structure, in order to continuously vary the rf power into the accelerator. Since the ferrite volume, which interacts with the rf power, is much larger than in the three-port circulator, the four-port circulator can be operated at a much higher power level. Yet, its insertion loss is much higher than a three-port one. The typical insertion loss and isolation values for four-port circulators are 0.3 and up to 30 dB, respectively.
OTHER MICROWAVE COMPONENTS WAVEGUIDE BENDS AND TWISTS, AND FLEXIBLE WAVEGUIDES The waveguide used to transport power from the microwave power source to the accelerator changes direction at several points. This is accomplished by bends and twists in the waveguide. Figure 5-13a shows two types of waveguide bends. They are designated as an E plane bend if the electric field changes direction and an H plane bend if the magnetic field changes direction. To avoid excessive reflections the cross section of the bend waveguide should be uniform and the radius of curvature should be large. With precision machining or casting
FIGURE 5-13 . Rectangular waveguide components. (a) E plane and H plane bends, (b)a waveguide 90" twist, and (c) a flexible waveguide.
techniques, it is possible to make low VSWR bends, with radius of curvature R r 1.5b for E plane bends and R r 1.5 a for H plane bends, where a and b are the width and height of the waveguide, respectively. Rectangular waveguides may be twisted to change the direction of polarization as shown in Figure 5-13b. Twists should be uniform and long in comparison to a wavelength to minimize reflections. For a 90" twist, the length should be longer than four times the guide wavelength. Figure 5-13c shows a flexible waveguide. Flexible waveguides, which can provide small bends and twists, are
98
CHAPTER 5. MICROWAVE POWER SOURCES AND SYSTEMS
often used to compensate for slight misalignment between rigid microwave components. A flexible waveguide also reduces mechanical stress on delicate rf windows of rf sources and accelerators.
Coupled
+
Ins Loss
DIRECTIONAL COUPLERS A directional coupler is a four-port device for transferring power from one transmission line to another in one direction, while isolating them from each other in the opposite direction. Figure 5-14 illustrates a schematic of a directional coupler and its symbols. A signal entering port 1 will travel to port 2 and a predetermined portion of this signal will appear at port 4, when all ports are matched. There will be no output at port 3. Similarly, if a signal travels from port 2 to port 1, output appears at port 3, but none at port 4. The coupling coefficient of a directional coupler is the ratio of the input power to the coupled output power, expressed in decibels. For instance, in a 20-dB coupler, 1 percent of power would appear at port 3 and 99 percent at port 2, if the input signal at port 1 is 100 percent. There are many types of directional couplers and no unique classification is possible. The type of directional coupler often used in linear accelerators is the 60 dB (which is 0.0001 percent of input power coupled) waveguide type as shown in Figure 5-15a and the low power 3-dB coaxial type quadrature hybrid coupler as shown in Figure 5-15b. A quadrature hybrid is a 3-dB directional coupler, capable of dividing an input signal into two mutually isolated quadrature phased outputs while maintaining isolation of the port 3 from the input, port 1. The term quadrature phased means that the two output
FIGURE 5-15 dB hybrid).
I
Coupled FIGURE 5-14 plers.
. Schematic diagram and symbols for directional cou-
signals are separated in time by 90" of rf period from one another.
SHUNT, SERIES, AND HYBRID TEES Figure 5-16a illustrates an H-plane tee, in which a waveguide is attached perpendicular to the narrow wall of a straight length of waveguide. This type of H-plane tee is often called a shunt tee, and the schematic view of the magnetic field distribution in an H-plane tee is illustrated in Figure 5-16b. If port 2 is shorted and
. Examples of directional couplers. (a) A 60-dB waveguide type coupler and (b)a quadrature hybrid power dividers (often called a 3-
OTHER MICROWAVE COMPONENTS
99
Figure 5-17b shows two types of hybrid tees that are often used in microwave circuitry. In the hybrid tee, the incident TElo mode waves in port 1 split equally and appear at ports 2 and 4 in phase. Since this electric field has even symmetry about the midplane, it cannot excite the TElo mode in port 3. Thus there is no coupling between ports 1 and 3. Similarly, the incident TE,, mode wave in port 3 will appear in ports 2 and 4 in equal magnitude but 180° out of phase. (a) Note that wavelength is shorter here
ROTARY JOINTS
FIGURE 9 1 6
. (a) H plane (shunt) tee and (b)magnetic field lines.
port 3 is terminated, the VSWR looking from port 1 to the load varies with the length I of the shorted waveguide associated with port 2. The equation of this relationship is given by
where s is the VSWR at port 1. In this way, the reflected power level can be varied continuously by varying the length of short position I. Varian high energy machines use this concept to vary the rf power level. Similarly, an E-plane tee, in which a waveguide is attached to the broad wall, is illustrated in Figure 5-17a with electric field distributions. This type of junction is called a series tee. These tees are used either to split the power or to match the impedance.
FIGURE 5-17
The gantry, on which the accelerator is mounted in high energy medical linacs, rotates around the patient. The rf source (klystron) and other rf components (circulator, shunt tee, etc.) remain fixed in stand type machines. Thus, it is necessary to have an rf joint that allows rotational motion but provides rf continuity. This can be accomplished by using a "choke joint," which allows relative rotation of two coaxial sections with electrical continuity (see Fig. 5-18a). The choke joint utilizes two quartenvave transmission lines of different impedances. The ratio of these impedances Zol and ZO2is given by
where r l = d2 - d l and r2 = d4 - d3. If one made r2 = lor,, an improvement in rf continuity by a factor of 100 over the method without these structure will be seen. This scheme is often used in microwave circuitry where a good rf connection is required without providing physical contact. Figure 5-18b shows a rotary joint using a coaxial line operating in the TEM mode as the transmission line mode, which is circularly symmetric and coupled to a rectangular waveguide. Suitable bearings are supplied outside of the junction so that the rotational motion is smooth. Figure 5-18c shows a cylindrical waveguide rotary joint where TM,, mode is used for propagation of microwaves within the circulator section.
. (a) E plane (series) tee and electric field lines and (b)hybrid tees.
100
CHAPTER 5. MICROWAVE POWER SOURCES AND SYSTEMS
Short Circuit
L FIGURE 5-18
"Across the Line and Guide Transition"
3:
I
Open Circuit
=o
Bearing Housing Dynamic Seal
4
- Choke joint and rf rotary joint. (a)Achoke joint, (b) a coaxial type rotary joint, and (c) a cylindrical waveguide type rotary joint.
WAVEGUIDE WINDOWS A waveguide window is used to separate the accelerator vacuum side from the gas pressurized (Freon or SF6) rf system side. Figure 5-19 illustrates a pill-box type window used in Varian accelerators. A high purity alumina disk (Al-300, which is 97% pure A1203) provides the ceramic
window. It is brazed into a thin cylindrical copper sleeve. The sleeve in turn is brazed to a stainless steel cylinder that is brazed to copper flanges, forming a pill-box cavity. The thin sleeve is provided to relieve stress caused by the thermal expansion differences between the ceramic window and the metal. The alumina disk is coated by sputtering with a thin
OTHER MICROWAVE COMPONENTS
7 h/4
101
choke
+E field (4 FIGURE 5-18
. (Continued)
layer of titanium nitride on the vacuum side after it is brazed to the sleeve assembly in order to reduce secondary electron emission and thus prevent multipactoring (a regenerative secondary electron emission process occumng at the ceramic window surface). Radio frequency window failure is very rare but could seriously damage the accelerator. The most
common failure is thermal failure with loss of vacuum integrity due to excessive localized heat produced in the window material. The sources of window heating are electrical breakdown, multipactoring and dielectric loss in the alumina, and resistive loss in the titanium coating. Therefore, the choice of window materials and techniques of coating and assembling of the window become very important factors in accelerator development and fabrication.
WATER LOADS
FIGURE 5-19
-
Cut away view of pill-box type rf window.
Loads are designed to absorb incident microwave power without reflection. Water loads are used for extremely high power termination. Two different types of waveguide water load are used in accelerators, namely, (a) water filled load and (b) water cooled dry load. Figure 5-20 shows the schematic cross section of these water loads. The water filled load is designed such that incident microwave power entering via the ceramic window is absorbed directly by the cooling water. The ceramic window behaves as a quarter wave matching section between the gas pressurized waveguide and the water filled section. Since the relative dielectric constant of water is about 80 at 3 GHz and E,, = 1, the dielectric constant of the window EW, is computed as E, = (E,,, X ~,,,,)1/2 = 9. A low loss dielectric material of dielectric
102
CHAPTER 5. MICROWAVE POWER SOURCES AND SYSTEMS
varying impedance mismatch with rotation, hence power reflection.
Nondirectional Coupler
Cooling Water lnlet & Outlet
Ceramic Window
AUTOMATIC FREQUENCY CONTROL The resonant frequency of an accelerator structure varies with temperature, input power level, beam loading, and other mechanical and electrical perturbations of the accelerating cavities. In order to keep the frequency of a microwave source, such as a klystron or magnetron, tuned to the accelerator resonance frequency, a frequency locking circuit called an automatic frequency control (AFC) is required. The methods are somewhat different for magnetron and klystron driven accelerators.
Cooling Water lnlet
LOW ENERGY (MAGNETRON) AUTOMATIC FREQUENCY CONTROL Aluminum
/ Housing
Figure 5-21 shows a schematic AFC circuit diagram for a low energy accelerator machine such as the Clinac 61100, which utilizes a magnetron as the microwave power source. The vernier mechanical tuner of the magnetron is sewoed to the resonant
Motor
t
Cooling Water Outlet
60-dB Magnetron Directional Coupler
Lossy Dielectric Material (Sic)
3 Port Circulator
Phase Accelerator Nondirectional
.
FIGURE 5-20 Cross section of types of water loads. (a) Filled with water, and (b) externally water cooled dry load.
Servo Drive Amplifer
constant 9 is alumina (A1203), which is also used in the rf window to provide an impedance match. The optimum thickness t of the water load window can be computed by
Water Load
3-dB Coupler
Attenuators
where Ag is the guide wavelength and represents a quarter wave length within the ceramic window. The water cooled dry load is constructed by mounting power absorbing lossy dielectric material (SIC) along the waveguide. Tapers are used to avoid reflection. For minimizing reflection, the taper length must be longer than one guide wavelength. The maximum average power handling capability for these loads is about 3 kW versus water loads that can handle over 10 kW. However, water loads cannot be used in he gantry since the air pocket within the cooling water creates
Detectors Differential Amplifier
Panel Meter
.
FIGURE 5-21 Block diagram of AFC circuit for low energy magnetron machines
103
AUTOMATIC FREQUENCY CONTROL
frequency of the accelerator using the reflected signal. The pulsed forward microwave power is supplied to the accelerator through port 1 of a three-port circulator. The reflected power from the accelerator travels back to the circulator via port 2 and out of the third port (lower) and, hence, via phase shifter to the water load. This variable phase shifter, often called a "phase wand," is provided forreflecting a small fraction of the reflected power from the accelerator back through the three-port circulator to the magnetron. This intentionally introduced reflected signal at a certain phase exerts a "frequency pulling effect" on the magnetron frequency towards the accelerator resonant frequency. This technique works well only when the frequency of the magnetron is within the one-half powerbandwidth of the acceleratorresonance. However, due to the relativelyhigh Qof the accelerator, the magnetron may have an initial frequency setting outside the bandwidth of the accelerator,since the mechanically tunable magnetron frequency range is about 10 MHz, while the accelerator resonant bandwidth is only about 0.5 MHz. Hence, another mechanism is needed to bring the magnetron frequency close to the acceleratorresonant frequency. In addition to the "frequency pulling effect" noted above, the microwave phase comparator circuit of Figure 5-21 is utilized to provide an AFC circuit that will track a relatively wide variation in the accelerator resonant frequency using the magnetron mechanical tuner. Departure of the source frequency from the accelerator resonant frequency can be sensed by comparing the phase of the incident power +i to the accelerator with the phase of reflected power 4, from the accelerator. A 3-dB hybrid coupler is provided for comparing this phase difference. Radio frequency signals from the two output ports are rectified by crystal detectors, and the resultant differential amplifier provides the input signals for the servopower amplifier that drives the servomotor until the correct magnetron frequency is reached. The frequency discriminating action of the phase comparator circuit, embodied in the 3-dB hybrid coupler, is shown in Figure 5-21. In the hybrid coupler, the incident power signal Vi, which is applied to the input port no. 1, is split and supplied to output terminals with equal amplitude but with 90°phase lag at the output port no. 2 with respect to the output port no. 4. Similarly, the reflected power signal V, applied to the second input terminal no. 3 is split equally and supplied to output terminals nos. 2 and 4 with a 90" phase lag at the output port no. 4 with respect to the output port no. 2. The resultant signals at output ports are rectified by crystal detector diodes to produce dc output voltages that are subtracted from each other and amplified with a differential amplifier.
Figure 5-22 shows the block diagram of the AFC system used for the Clinac 1800. Here the function of the oscillator (rf driver) and the amplifier are separated. The frequency of the rf driver can be set precisely at a desired accelerator resonance, independent of the load condition and the frequency tuning range can be set in a relatively narrow range, such as 1 MHz. The center frequency of the rf driver is set nominally at 2856 MHz. The forward and the reflected signals are monitored at the accelerator guide through a 60-dB bidirectional coupler. Frequency discrimination can be obtained in a manner similar to that described for the low energy machine. In this case, the delay line used for the low energy machine is not necessary since the next nearest resonance will not appear within the rf driver tuning range. Figure 5-23 shows how the phase and amplitude of reflected signals vary with frequency for the case of the accelerator overcoupled to the input waveguide, where the accelerator resonant frequency is fo. The frequency discriminating action of the AFC circuit can be described by the vector diagram of Figure 5-24, which shows the relationship of the klystron frequency f versus rf signals at the hybrid coupler output port and, hence, the differential amplifier output signal. However, this relationship applies only for the condition where the transmission path length for the incident wave Vi and the reflected wave V, are equal. Normally, the path length for the reflected wave V, is greater than the path length for the incident wave Vi, whereby the frequency discriminating output characteristic for the phase comparator would be skewed. Accordingly, a long line section of transmission line (often called delay line) is
,
Pads Are Typically 5 dB (but may vary)
60 dB
Direction Cou~ler
Shifter
30
Hybrid
AFC Tuning to RF Driver AFC
HIGH ENERGY (KLYSTRON) AUTOMATIC FREQUENCY CONTROL In high energy accelerators, in which klystron amplifiers are used as therf power sources,the AFC circuit is somewhat different from that in low energy magnetron powered accelerators.
2.2KQ
I317 on Program Board
FIGURE 5-22 machine.
. Block diagram of AFC circuit for high energy klystron
104
CHAPTER 5. MICROWAVE POWER SOURCES AND SYSTEMS
provided in the path between the forward output port of the 60-dB directional coupler and the input port of the 3-dB hybrid coupler, as shown in Figure 5-21.
Freq.
* Freq.
0. A
2
-
FIGURE 5-23 Voltage and phase versus frequency for frequency discriminating action of AFC circuit.
Frequency
Output Port
Output Port
No. 1
No. 2
Differential Amplitude
FIGURE 5-24 . Vector relationships at hybrid output port and resulting differential amplitude at various frequencies.
REFERENCES 1. Granastein VL and Alexeff I: High power microwave sources. Artech House, Inc., 1987. 2. Veley VF: Modem microwave technology. Prentice-Hall, Inc., 1987. 3. Gilmour AS: Microwave tubes. Artech House, Inc., 1986. 4. Linkhart DK: Microwave circulator design. Artech House, Inc., 1989. 5. Hinkel K: Magnetrons. New York, John F. Rider Publisher, Inc., 1961. 6. Harvey AF: Microwave Engineering. Academic Press, 1963. 7. Rizzi PA: Microwave Engineering, passive circuits. PrenticeHall, 1988. 8. Cheung WS and Frederic HL: Microwaves made simple: principles and applications. Artech House, Inc., 1985. 10. Collins, GB: Microwave Magnetrons. McGraw-Hill Book Co., Inc., 1948.
Pulse Modulators and Auxiliary Systems
PULSE MODULATORS The pulse modulator provides a pulsed voltage or current waveform to power microwave sources (magnetrons or klystrons). Figure 6-1 shows a basic modulator configuration. The electrical energy, supplied by a power supply circuit, is stored in an energy storage circuit. An isolating circuit prevents flow of the stored energy back to the power supply. A switch circuit provides for the discharge of energy in pulses into the load. Figure 6-2 shows a simplified circuit of a line type modulator. The name, line type, came from the similarity of the behavior of the energy storage circuit to that of a transmission line having a characteristic impedance &.Such a line is charged to voltage V and then discharged to a load with an impedance typically equal to that of the transmission line and with a wave propagation velocity v. As shown in Figure 3-4, transmission lines can be represented as a network of lumped capacitive and inductive elements. Such a network (see Figure 6-3) is called a pulse forming network (PFN). The PFN serves a dual purpose of storing the energy required for a single pulse and of discharging this energy into a load in the form of a pulse of specified shape. The PFN of Figure 6-3 consists of N equal-valued capacitors, Cl..C6,and a continuously wound, tapped coil representing Nequal-valued inductances L1..L6,and whose physical dimensions are chosen such that it provides the proper mutual coupling at each mesh.
Isolating Circuit
1 FIGURE 6-1
Switch Circuit
Circuit
. Block diagram of basic modulator.
Load
i
The PFN capacitors are charged through a resonant charging choke to 2V, where V is the power supply voltage. When the switch S , is closed, the charge stored in C1 starts to discharge through L, into the load R,. The characteristic impedance ofthe PFN is designed to be equal to the load impedance. When the voltage across the load reaches V, C1 will stop discharging, and C2willbegin to discharge through and L,, and so on, through C&. This sequence of events results in a rectangular current pulse of energy being supplied to the load with a duration that is twice the transmission time of the PFN. In essence, a negative wave of amplitude Vl2 flows from the load end to the charge end of the PFN, reflects and flows back to the load end, discharging the stored energy in the capacitors fully. The important relationships of the PFN are
where Zo = characteristic impedance of PFN
T = pulse width L, = total network inductance = NL C, = total network capacitance = NC L = inductance per section C = capacitance per section N = number of sections
The PFN used in the CL-1800 modulator has six sections and a characteristic impedance of 11.4 In order to obtain a 5.8-ps pulse width, one can solve eqs. (6-1) and (6-2) in terms of L and C, and obtain L = 36.5 pH and C = 0.28 pF. The modulator must supply the klystron or magnetron with a pulse of precise amplitude. One method of PFN charge regulation involves diverting the discharge of stored energy in the charging choke, by abruptly reducing the "Q" of the charging circuit. This is called deQing. Figure 6-4 shows the resonant charging system with De-Q. By sensing the PFN charging
a.
106
CHAPTER 6. PULSE MODULATORS AND AUXILIARY SYSTEMS
30 High Voltage PWR Supply
Clipper Current Fault Monitor System
Circuit = 1000:l Ratio
-
-
1 0 kV
Current Toroid
Ll O
n
-
\ r
HVPS Current Monit0;Circuit --R2
I
L
I
HVPS O/C Fault Monitor System
----
I
-1
L-----J
Magnetron Equivalent Circuit --------
o
_I
Despiking Network
I
End Clipper Circuit
Pulse XFMR Turns Ratio = 1:4
I
I I
I !
'VVL
o
DeQing Switch Trig. Gen. System
r---1r.--
--A
-
I I
I----- J I I
I I I
vs=
I
I
. Simplified circuit of line type modulator.
FIGURE 6-2
Charging Choke U
PFN
-
---Cl
--c2
----
----
c3
c4
I
Switch
---c5
---C6
---
4
:- - - - - - - - - - - - - - - - - -
A
FIGURE 6-4 Schematic of PFN resonant charging with regulation of charge by De-Q network. -4
FIGURE 6-3
. Pulse forming network.
Voltage Sensor
107
VACUUM SYSTEMS
voltage and closing switch S2,the charging process will be terminated by dumping the residual energy stored in the charging choke into the resistor R,.
VACUUM SYSTEMS
-
An accelerator must be kept at high vacuum (pressure of 10-7 torr level) (1 torr = 1 mmHg, 1 atm = 760 ton) to prevent electrical breakdown in the residual gas forthe high electromagnetic fields used to accelerateelectrons. Figure 6-5 shows a schematic view of the Clinac 1800vacuum system. There are three Vacion pump subsystems; two 2-Lls diode Vacion pump systems for klystron and electron gun, and one 20-Lls triode Vacion pump for the accelerator and the bend magnet vacuum chamber. Figure 6-6 shows a schematic of two Vacion pumps: (a) a diode pump and (b) a triode pump. The external permanent magnets are 0.1-0.3-tesla strength and cathode voltages of approximately 5 kV are applied. The electric field traps the electrons in a potential well between that of the two cathodes and the axial magnetic field forces the electrons into circular orbits, which prevent their reaching the anode quickly. The combination of electric and magnetic fields causes the electrons to travel long distances in oscillating spiral paths before colliding with the anode and results in a high probability of ionizing collisions with residual gas molecules. The ions produced in these collisions are accelerated toward the cathodes,
where they collide, sputter away the titanium cathode, and release secondary electrons that in turn are accelerated by the electric field. The mechanism of pumping is dependent on the nature of the gas being pumped and is based on one or more of the following mechanisms. 1. Trapping of electrons in orbits by a magnetic field. 2. Ionization of gas by collision with electrons. 3. Sputtering of titanium by ion bombardment. 4. Gettering of active gases by titanium. 5. Diffusion of hydrogen and helium into titanium. 6. Dissociation of complex molecules into simple ones for easy pumping. Organic gases, active gases, hydrogen, and inert gases are pumped in different ways. Organic gases are easily pumped by adsorption and precipitation after being dissociated by electron bombardment. Active gases such as oxygen, carbon monoxide, and nitrogen are pumped by reaction with titanium, which is sputtered on the anode surfaces, and by ion burial in the cathode. Hydrogen is initially pumped by ion burial and neutral adsorption and diffuses into the bulk of the titanium forming a hydride. Inert gases are not pumped efficiently as active gases in a diode pump. Argon, in particular, suffers from a pumping instability. In the triode pump, as shown in Figure 6-6b, argon ions are neutralized by glancing collisions with the sputter cathode, impact the pump wall, and are covered with sputtered
I
Stand
Gantly
I I
I I 1 I I I I I I
Klystron
Diode Vacion Pump
I
Gun
Diode Vacion Pump
I +3.2 kV
Triode Vacion Pump A1 9
A13
A
h +3.2 kV
I I
I I FIGURE 6-5
. Block diagram of vacuum system for Clinac 1800.
Bend Magnet
Accelerator
Vacion PIS
-5 kV
108
CHAPTER 6. PULSE MODULATORS AND AUXEIARY SYSTEMS
Ti?
Control Unit
I
7 7
--
control Unit
Magnet
Multi-Cell Anode Pump Wall Forms Third Electrode
I
Sputter Cathodes, Titanium Vanes
Multi-Cell Anodes
(3)
FIGURE 6-6
. VacIon pumps. (a) A diode type and (b)a triode type. Target Mechanism Gun
2-Liter Vacion
Accelerator Guide
A13
Diode Type
IQ
Nitrogen Let-Up Valve
I
Magnet Vacuum
Ceramic Vacuum Window
Bellows Connection
I
Over Research Port
Beryllium Window
i
Useful Down to 10 torr
.
Valve
20-Liter Vacion
torr
JlC Gauge Useful Down To torr
.......................
Coaxial Trap
Pump-Down Butterfly Valve
FIGURE 6-7
. Vacuum components in the gantry of a Clinac 1800.
1.75 CFM (Stops Pumping ~ at 1o - torr)
Sargent-Welch Mod. 8805
109
VACUUM SYSTEMS
titanium. The pump wall surface in the triode pump operates at the anode potential and collects low energy ions that could not sputter. The lifetime of a diode pump is a function of the time necessary to sputter through the cathodes. A typical value is 5000 h at 10-5 torr or 50,000 h at 10-6 torr. Vacuum requirements are more stringent in the klystron than in the accelerator since the klystron oxide cathode is highly susceptible to poisoning by atmospheric gases and vapors. The klystron pressure (5 X 10-8 torr) is an order of magnitude lower than the accelerator pressure and is separated from it by rf windows. Figure 6-7 shows the vacuum components in the gantry. The accelerator is initially evacuated by a vane type roughing pump through a coaxial oil trap. The trap is filled with copper wool to provide a large surface area to prevent oil leaving the pump from contaminating the accelerator. While the roughing pump is evacuating the system, the degree of vacuum is monitored by a thermocouple gauge. When the roughing pump has evacuated 99.9999 percent of the air, the vacuum pressure in the accelerator is down to 10-3 torr. Then the main valve is closed and the 20-Lls Vacion pump is turned on, pulling the accelerator down to a pressure of 10-7 torr. Figure 6-8 shows the relationship of Vacion current versus
10 pA
pressure for various Vacion pumps. Since the conductance of the accelerator for molecular flow is so low, the pressure level in the accelerator is higher than that shown by the pump curve. The basic vacuum system relationships are Q=CXAP=SXP=DXA
Q = throughput (gas flow) in torr times liters per second C = conductance in liters per second AP = pressure difference in torr D = surface outgassing rate in tom x liters per second x square centimeters A = total internal surface area in square centimeters S = pumping speed in liters per second
For example, consider the SLAC constant impedance TW structure (see Figure 6-9) Operating mode 21~13 Number of cavities 50 Accelerator length 1.75 m Inside diameter 3.247 in. Disk hole diameter 0.890 in.
100 pA
1 mA
Pump Current
FIGURE 6-8
(6-3)
where
. Relationship of VacIon current versus pressure for various pumps.
10 mA
100
110
CHAPTER 6. PULSE MODULATORS AND AUXILIARY SYSTEMS
The cavity conductance Cc is calculated as follows:
cc= 453.9 U s Then the accelerator total conductance Camwill be
where N is the number of cavities. Since N 0.67 Us. 2. Accelerator Guide
Waveguide
I
1 ft
= 50,
C ~ C=C
The conductance of coupling aperture assuming infinitely thin wall can be computed from
F L
where a and bare aperture height and width in centimeters. Thus,
RF Window
I
I
Vacion Pump
3.
The conductance of the 12-in. rectangular waveguide is calculated from
Pumpout Tubing
FIGURE 6-9 . Vacuum system of a SLAC TW accelerator structure, for computation of conductances and pumping speed.
Disk thickness 0.230 in. Period length 1.378 in. Matching iris aperture 1.014 X 1.34 in. Waveguide dimension 2.84 X 1.34 in. Pump tubing diameter 1.115 in. Pump speed 20 U s
where a and b are dimensions of the waveguide (cm), L is the length (cm), and K is a factor depending on the ratio of bla. For a = 7.214 and b = 3.404, Kis about 1.16. Hence c w = 57.5 L/s. 4.
The conductance of the 12-in. length pump tubing Cp is calculated from eqs. (6-4) and (6-5), hence or = 0.85, Cp = 7.6 Us. 5. The total conductance CT is computed by
Neglecting the input and output beam holes and coupling iris to the load, the pressure ratio between the end cavity and the pump can be computed as follows: 1. The conductance of accelerator vacuum components, C, can be calculated as follows:
where D is the diameter of the tube (cm), L is the length of the tube (cm), and or is the Causing factor depending on the ratio of LID and given by
hence CT = 0.61 Us. Since Ppump
'pump =
CT (Pacc - ppUmp)
The pressure at the far end of the accelerator is 34 times higher than the pressure at the pump for this example.
WATER COOLING SYSTEM
Therefore, the aperture conductance Ca is calculated as follows:
The water cooling system cools critical components such as accelerator guide structures, rf power source, and electromagnets in the drive stand and gantry, keeping them at a relatively constant temperature. The cooling system contains a submersible water pump enclosed in an 18-gal tank, a heat exchanger
111
MISCELLANEOUS SYSTEMS
Heat Exchanger
Y City Water Supply
-
--I(t
Water Return Path
II l t
I :r:
Stand Components
Gantry Components
I
1
cM.2hb cu a
5
a 18 Gal Cap
Rotary Joint
-
E
5 (I,
Submersible pump in 18 gallon water tank pumps demineralized water through heat exchanger, Then through components in stand and gantry that require stable operating temperatures
$ -IN 7
FIGURE 6-10
. Block diagram of basic water cooling system.
connected to the hospital city water supply, return manifolds, water hoses, temperature and pressure gauges, and temperature control circuitry (see Fig. 6-10). For the Clinac 2100, assemblies and components in the gantry and stand are cooled by circulating water; in the gantry: accelerator guide, bend magnet, accelerator solenoid, primary collimator, energy slit, and target; and in the stand: klystron, klystron solenoid, circulator, rf driver, rf load, and pulse transformer. The water, at a temperature of 40 5°C and at a pressure of 50psi, enters the gantry from the stand through a rotary joint, cools the components in the stand, and returns to the stand through another rotary joint shown in Figure 6-11. If the volume of water flowing through any of the flow interlock switches falls below the values indicated the switch will open, transmitting a flow fault signal to the console and where appropriate, shut down the associated systems. The heat exchanger assembly isolates the city water from the accelerator water closed-loop circulatory system as shown in Figure 6-10 while allowing heat to be exchanged between the two water circuits. The city water is connected by hoses and filtered before reaching the heat exchanger. A pressure gauge monitors the incoming city water, and controls the flow by an electromechanical control unit, valve, and temperature sensor
+
in the closed-loop water circuit. A submersible pump in the 18-gal tank circulates distilled water at a nominal pressure of 50 psig throughout the closed-loop water circuit. Interlocks monitor and protect the water level and tank temperature. The interlocks also activate the pump fault lamp on the console. The water is kept pure by a filter between the pump and heat exchanger and also by a demineralizer cartridge in the rf driver branch of the closed-loop system. The temperature of the cooled water is monitored at the output of the heat exchanger. When the temperature deviates, the electromechanical temperature controller sends a signal to the control valve, changing the flow rate of the city water; thereby correcting the amount of heat transferred.
MISCELLANEOUS SYSTEMS GAS DIELECTRIC SYSTEM Pressurized gases such as SF6 or Freon 12 are used as a dielectric for the waveguide system between the rf source and the accelerator. A gas dielectric assembly supplies sulfur hexa-
112
CHAPTER 6. PULSE MODULATORS AND AUXILIARY SYSTEMS
7 A
FIGURE 6-11
. View of cooling water supply to various gantry components.
fluoride (SF6) in gaseous form under pressure to the waveguide, between the klystron window and the accelerator guide window. Freon 12 gas is normally used in a magnetron system. The gas acts as a high dielectric that minimizes arcing in the waveguide. Figure 6-12 compares two pressurized gases with air for peak power handling capability. This figure shows that SF6gas at a pressure of 25 psi operates at 10times higher power than air at atmospheric pressure. The gas dielectric assembly is located on the side of the drive stand. It consists of an SF6 tank, a single-stage regulator and shut-off valves, copper tubing, drier filter, solenoid valve, relief valve, diaphragm valve, pressure switch, and terminal board. The SF6gas is stored in a metal tank at a pressure of 310 psig. The gas is supplied through copper tubing to the gas dielectric assembly, where it is dried through a filter and regulated to a normal pressure of 32 psig by a solenoid activated valve. The pressure is controlled by two pressure switches. If the gas pressure drops to less than 30 psig, the no. 1 switch opens to let in more gas until the pressure is 32 psig, then closes. This continues until the gas bottle is depleted.
Relative Power Characteristics
100 1
Absolute Pressure (lbf/in2a)
.
FIGURE 6-12 Relative peak power breakdown threshold versus pressure for nitrogen, air, and SF6.
MISCELLANEOUS SYSTEMS
When this occurs, the no. 2 switch opens and activates the gas pressure interlock circuitry, causing beam termination and the gas lamp on the console to light.
113
sure will cause the air lamp on the console to light and cause an automatic shut-down of the system.
REFERENCES PNEUMATIC SYSTEM This system provides pressurized air for air driven mechanisms. Pressurized air is used for various purposes; (a) to move the target, (b) to operate the locking pin plungers on the carrousel, (c) to operate the plungers on the shunt tee, and (d) to move the energy switch. The air pressure is controlled by an air regulator assembly, and set between 45 and 50 psig. The application of air pressure to all the drive mechanisms is turned on and off by electrically operated air control solenoids. Abnormal air pres-
1. O'Hanlon JF: A user's guide to vacuum technology. John Wiley & Sons, 1980. 2. Guthrie A: Vacuum equipment and techniques. McGraw-Hill Book Co., 1949. 3. Hablanian MH: High-vacuum technology, a practical guide. Marcel Dekker, Inc., 1990. 4. Guthrie A: Vacuum Technology. John Wiley & Sons, Inc., 1963. 5. Glasoe GN and Lebacquz JV: Pulse generators. McGraw-Hill Book Co., Inc., 1948.
C H A P T E R
Beam Optics of Magnet Systems
OVERVIEW This chapter is an introduction to the subject of how magnets are used to confine, deflect, and focus the electron beam in medical accelerators. The emphasis is on the motion of electron rays and bundles of such rays through individual types of magnets and through systems of magnets. This first section discusses the need for a bend magnet in high energy medical accelerators and the need for its beam orbit height above the x-ray target to be limited, thereby restricting the choice of magnet geometries. The second section (Electron Motion in Magnetic Fields) presents the basic physics of electron motion in free space and in the magnetic forces of dipole, quadrupole, solenoidal, and steering magnets. The third section (Beam Emittance) discusses a very useful concept for treating bundles of electron rays, namely, beam emittance. The next two sections (Nonachromatic Bend Magnet Systems and Achromatic Bend Magnet Systems) describe a number of different bend magnet types that have been used in medical accelerators from the point of view of their ability to faithfully transport a beam of significant energy spread from the output of the accelerator structure to the x-ray target, that is, their degree of achromaticity.
STRAIGHT AHEAD LINACS Isocentric linacs of straight-ahead beam design are produced with x-ray energies of 4 or 6 MV without the use of magnets. The SW accelerator structures in these machines are typically 25-35 cm long, so short that a beam diameter of about 3 mm is achieved at the x-ray target without the use of a solenoid or magnetic lens to confine and focus the beam. Such machines employ an accelerator structure of high shunt impedance to keep it short, a short gun, and a minimal length x-ray target structure in order to limit the isocenter height with 360" gantry rotation. The accelerator structure must be able to sustain a high electric field, having an energy gradient of approximately 20 MeVIm.
BENT BEAM LINACS Isocentric linacs of bent beam design achieve acceptable isocenter height without severe constraints on the length of the gun, accelerator structure, and x-ray target region. Bent beam design is used by choice in some 4 and 6-MV x-ray energy machines and it becomes essential for machines of higher x-ray energy. The accelerator structure in such machines is approximately horizontal, typically 1-2.5 m long, with solenoid andlor magnetic lenses to maintain small beam diameter over this length. A magnet system bends the beam through a net angle of approximately 90"-270" onto the x-ray target, electron output window, electron scattering foils, or electron scanning magnet axis. Such an array of magnets and drift spaces is termed the beam transport system. In medical linacs it is comprised primarily of the bend magnet. The characteristics of linac electron beams and their motion in various types of magnets are discussed in the following sections.
LINAC BEAM CHARACTERISTICS The accelerated electron beam is comprised of an array of electrons each of which differs in energy (hence momentum) and in radial displacement and radial divergence (angle) from the electron ray representing central momentum and central trajectory. Energy slits and collimators are used to limit the energy spectrum width and the beam cross section and angular divergence, which are accepted by the transport system for delivery to its output. Typical beam transport acceptance values for medical linacs are in the ranges of 3 to 10 percent energy spread full width, 2 to 4-mm beam diameter, and 2 1 to 5-mrad (milliradian) angular divergence from axis.
EFFECT OF MAGNET SYSTEM CHOICE ON ISOCENTER HEIGHT For convenience in setting up the patient, the isocenter should not be too high above the floor. Figure 1-20 shows a radiation
7
116
CHAPTER 7. BEAM OPTICS OFMAGNET SYSTEMS
technologist of typical height. Assuming a 30-cm thick patient with upper surface at the radiation technologist's eye level of 150 cm, the center of the patient would be 135 cm above the floor. It is desirable to have an isocenter height of less than this 135 cm, so this requires use of a bend magnet type with limited beam orbit dimensions. With 360" isocentric gantry aimed upward and 100-cm source axis distance, there is a space of less than 35 cm between the x-ray target and the floor for the following: Rotation clearance Radiation head enclosure thickness Supplementary radiation shielding Magnet return yoke Yoke to pole edge spacing Pole edge to beam trajectory margin Beam trajectory height h above the x-ray target. One example of the space required by these items is illustrated in Figure 7-1. In order to limit beam aberrations due to magnet saturation effects, practical bend magnets are usually not operated above about 1.7 T (Tesla) (17 kG), which corresponds to a radius of curvature of 5 cm at 25 MeV. To permit comparison of orbit designs on a common basis for the various types of bend magnet systems used in medical electron accelerators, the ratio of the trajectory height h above the x-ray target, to the radius of curvature p can be used. If all systems employed the same
radius of curvature of p = 5 cm, the value of h would be between 9 and 18 cm, a range of 9 cm (3 112 in). In the magnet systems of Figures 7-19, 7-22, and 7-23, this value of h could have been reduced by about 5 cm by locating the x-ray target near the intersection of the exit and entrance beam trajectories, as is done in the magnet system of Figure 7-21.
ELECTRON MOTION IN MAGNETIC FIELDS ELECTRON MOMENTUM The product mv is termed momentump, where m is the relativistic mass. The ratio P of electron velocity v to the velocity of light c is a function of kinetic energy T. Defining the electron rest energy as Wo = rnoc2, the total energy W = Wo T and their ratio WIW, as y:
+
The ratio of moving electron rest mass m to electron rest mass mo is
The momentum is
It is convenient to state electron momentump* in units of moc.
The electron rest energy, Wo = moc2is 0.510985 MeV. For example, at T = 10 MeV,
P = 0.998818, y = 20.5700, p = 20.5457 m,c m0 = 9.1085 X g, c = 2.997929 X 101Ocm/s, e = 1.60207 X Coulomb
ELECTRON MOTION IN THE DIPOLE MAGNETIC FIELD FIGURE 7-1 . Example of bend magnet cross section in radiation head, showing limited total space and typical space for individual items.
Figure 7-2 shows the force F, experienced by an electron moving in a magnetic field. The direction of the force is at a right angle to the direction of the magnetic field and at a right
ELECTRON MOTION IN MAGNET FIELDS
-
FIGURE 7-2 Direction of force, F = mv2/r, on an electron moving in the median plane of a magnetic field.
angle to the direction of the electron motion, deflecting the electron on a circular trajectory of radius p. The resulting centrepital force F = ntv21p is balanced by the centrifugal force Bev. The term Bev is called magnetic rigidity. Equating these two forces give Bp = rnvle = ple
(7-5)
Substituting from eq. (7-4),
For example, at T = 10 Mev, Bp = 35,019 G cm, and B = 7004 G for p = 5-cm radius of curvature. For beam energies typical of medical accelerators eq. (7-6) can be approximated by the following, with B in kilogauss, p in centimeters, T i n million electron volts.
Figure 7-3 shows that the radial bending of electron rays that are parallel to each other on entry normal to a uniform field magnet sector tends to converge these rays in the median plane toward a focal point for the geometry illustrated. It also shows the dispersion of these rays and their foci due to their different momenta.
ELECTRON MOTION IN THE FRINGE FIELD AT THE EDGE OF THE DIPOLE MAGNET The magnetic lines of force in the gap between the poles tend to repel each other, bulging out of the entrance and exit faces of a
117
FIGURE 7-3 . Dispersion of foci due to momentum spread in a dipole sector magnet.
bend magnet. In these fringe field regions, the magnetic field has vector components normal to the entrance and exit faces, which act on electron rays that are displaced in they direction above and below the magnet median plane and that are at an angle to these faces other than 90" (see Fig. 7-4). If the plane of the input face is rotated away from the center of curvature of the bending beam, there is a component of the force that pulls the y displaced rays toward the median plane during passage through the fringe field. Conversely, a ray in the median plane and displaced in its x direction toward (or away from) the center of curvature from the central ray passes through an extra (orlesser) region of magnetic field due to this entrance face rotation and is deflected away from the central ray. Thus, rotating the entrance or exit face of a bend magnet produces the same effect as a quadn~pole(see the next section) magnet at that point, causing beam convergence in one plane and beam divergence in the other plane. The process of bending results in initially parallel rays converging toward each other in the median plane as shown in Figure 7-3. By rotating the input and output pole faces a parallel beam can thus be focused in both radial and transverse planes to a point beyond the bend magnet, as illustrated in Figure 7-4b and 7-5b and discussed by Engels and Livingood.?-5 In the midplane of the magnet the fringe field starts rising from zero at some point outside the magnet and reaches full value (or nearly so) a short distance inside the pole gap (see Figure 7-4c). For simplicity in transport calculations, a "hard edge" approximation is made in which the field is assumed to step from zero to full value at a point a distance d from the pole edge. The ratio of d to the total pole gap g is defined as the constant k, which typically has a value between 0.4 and 0.7. The value of k can be reduced significantly by use of a field clamp, as illustrated in Figure 7-4d. The field clamp shunts much of the fringe field around the beam. This is especially useful in magnets having a total gap that is a significant percentage of the beam path length within the poles. Since bend magnets for medical linacs are necessarily compact, gaplpath length ratios are often fairly large and focusing properties are
Elevation Cross-section
Median Plane
k = dlg
Hard Edge
Hard Edge
z
FIGURE 7-4 Fringe field. (a) In the radial plane offset by y from the median plane: Direction of force on an electron moving in the fringe field of a rotated input pole face of a dipole sector magnet. (b) In the transverse plane: Effect of input pole face rotation on position of focus (from Ref. 4). (c) Fringe field distribution and hard-edge approximation without field clamp. (d) With field clamp.
119
ELECTRON MOTION IN MAGNET FIELDS
netic flux lines, and the lines of force orthogonal to them, are rectangular hyperbolas. For the pole polarities shown and for electrons coming out of the paper in Figure 7-6a:along the x axis, the magnetic field is in the y direction, the force on an electron is in the x direction, and both field and force increase linearly with x from zero at the axis; along the y axis, the magnetic field is in the x direction, the force is in they direction, both field and force increase linearly with y from zero at the axis. By = gx, B, = gy, where g is the magnetic field gradient in units of change in gauss per centimeter. For the polarity shown and for electrons coming out of the paper the magnet is focusing in the y plane (Fig 7-6b), defocusing in the x plane (Fig 7-&). To first order, the motion in the focusing plane is described by sinusoidal functions; the motion in the defocusing plane by hyperbolic sinusoidal functions: Y(Z) = y1 cos wz + y; w-' sin o z Y'(z) = -yl sin o z + y; cos a z x(z) = X, C O S ~o z x,' a - I sinh oz ~ ' ( 2 )= x , o sinh o z +x cosh Wz
+
(7-8)
where x l , y l ~ ' land y f l are the initial positions and angles relative to the axis and z is the distance along the axis. the B0 parameter w is the ratio of magnetic field gradient - and the a
product Bp of the magnetic field and radius of curvature for the particle. Hence,
where Bo is magnetic field magnitude at the surface of the pole at minimum distance a from the axis. For example, Bp = 35,019 G cm for a 10 MeV kinetic energy electron. Let Bo = 5 kg at the pole at a = 0.5 cm from 7r
FIGURE 7-5 . Effect of input pole face rotation on position of focus in median plane. (a) Input face normal to beam and (b) input face rotated away from beam.
improved by use of field clamps. The positioning of such field clamps also provides an additional parameter for tuning of the magnet focusing properties.
ELECTRON MOTION IN THE QUADRUPOLE MAGNETIC FIELD Figure 7-6a shows the shape of the magnetic field and magnetic equipotentials produced by a quadrupole magnet. The mag-
the axis. Then wz would equal- at 2.75 cm. Thus, a quadrupole 4 doublet suitable for focusing 10-MeV electrons could be installed in an axial length of less than 10 cm. A single quadrupole magnet (quadrupole singlet) converts an initially circular cross-section beam into an elliptical cross section. Usually, quadrupole magnets are used in pairs (quadrupole doublet) or in threes (quadrupole triplet), with pole orientations such that in a given plane defocusing in one magnet is followed by focusing in the next magnet, permitting stigmatic focusing (circular output beam cross section for circular input beam cross section). Ray paths in yz and xz planes are shown in Figure 7-6b and c for parallel to point focusing in a quadrupole doublet. The ratio of imagelobject size is termed the magnification M. The spread in divergence angles in the image plane is proportional to 11M (see Fig.7-7). That is, increased image size corresponds to reduced spread of divergence angles of rays from a point at the object. A quadrupole singlet is sometimes used before a bend
120
CHAPTER 7. BEAM OPTICS OFMAGNET SYSTEMS
FIGURE 7-7
. Formation of paraxial image with magnificationM.
magnet to correct for astigmatism in the bend magnet. Quadrupole doublets and triplets may be used to transport a beam over a considerable drift distance between the output of an accelerator structure and the region of application of the beam. An example might be one accelerator feeding beams to two gantries in separate treatment rooms. Where space permits, the triplet is superior to the doublet since it provides more symmetric focusing of the beam and permits independent control of focusing in the two transverse planes.
ELECTRON MOTION IN THE MAGNETIC FIELD OF A SOLENOID
FIGURE 7-6 . Quadrupole magnet. (a) Magnet equipotentials (solid curves,xy = constant); lines of force (dashed curves, y2 - x2 = constant). Forces on an electron coming out of the paper. Single arrows indicate magnetic field direction; double arrows indicate force direction. (b)Motion of an electron in transverse yz plane and (c) in transverse xz plane of quadrupole doublet and triplet.
The word solenoid is derived from Greek words meaning like a pipe or channel. Its magnetic field confines the diameter of flow of the beam electrons. A solenoid magnet is often used over at least a portion of the length of the accelerator structure in bent beam medical electron accelerators. It produces a field that is parallel to the accelerator axis. One purpose is to limit expansion of the beam diameter due to its spread in divergence angles at injection so that it will pass through the accelerating cavity apertures. A second purpose is to compensate for radial defocusing forces of the accelerating microwave field in a portion of its cycle in the early part of the accelerator structure where the electron beam is less relativistic. Electrons that pass through the midplane of an accelerator cavity gap, while the accelerating electric component of the electromagnetic field is rising (i.e., ahead of the crest of accelerating field), experience a longitudinal bunching force (early electrons accelerated less than late electrons) but experience a net radial defocusing force. The electrons experience less convergent electric force on entering the gap than divergent electric'force on leaving the gap. This electric radial outward force is partially compensated by a magnetic radial inward force due to the product of the azimuthal magnetic component of the electromagnetic field and the longitudinal velocity p,c of the electron. The crest of the magnetic component is 90" ahead of the electric component of the electromagnetic field. This net difference of electric and magnetic radial forces is thus dependent on the phase of the electron and it also varies as 1 - bePo,where Po is the phase velocity of the electromagnetic field. In medical accelerators the elec-
121
ELECTRON MOTION IN MAGNET FIELDS
tron velocity and the electromagnetic field phase velocity rapidly approach the velocity of light. Thus, the radial defocusing force drops rapidly as the electron gains energy, so external means of focusing to compensate is needed primarily in low beam energy regions of the accelerator guide. For &,Po the required solenoidal magnetic field B, to just compensate for the net electromagnetic radial force, is shown by Chodorow et al.14 [their eq. (5.23)] to be
For example, let EeAlm0c2 = 4, corresponding to 20 MeVIm at 10 cm wavelength, sin 0 = 0.5 (electron 30" ahead of the electric field crest, 60" behind the magnetic field crest), and p = 0.70, corresponding to 0.2-MeV beam energy. Then B = 824 Oe. For these same conditions, the required value of solenoidal field drops to 414 Oe at 1 MeV, 314 Oe at 2 MeV, 192 Oersted (Oe) at 6 MeV. Of course, a larger value than this just compensating value of B can be used. In high energy and, especially, dual x-ray energy medical accelerators, the solenoid may extend over most of the accelerator structure length and produce a uniform field of about 1000 Oe. In low energy accelerators, especially those that do not employ a bend magnet, a solenoid may not be used because the accelerator structure length is so short that radial momentumof the electron does not have time to produce excessive radial displacement at the x-ray target. Also, radial outward momentum produced in the first accelerator structure cavity can be partially compensated by injecting the electrons from the gun into this cavity with radial inward momentum. The lines of magnetic flux are parallel to its axis z inside the solenoid and flare out at its ends (usually ending on the
aperture of an enclosing iron cylinder), as shown in Figure 7-8. In this end field, the integral of the radial component of magnet field over a path parallel to z and displaced from the axis by r is equal to one-half the integral of the z component of the field inside the solenoid over this displacement distance r. Thus, the azimuthal acceleration given an electron by this fringe field is proportional to the displacement of the electron from the axis. The resulting azimuthal motion in the axial field produces a radial restoring force. All electrons entering parallel to the axis form helices tangent to the axis, the helix diameter equaling the displacement from the axis of the entering electron. Electrons entering with a divergence angle follow correspondingly larger diameter helical paths, the radial component of momentum due to divergence adding vectorially to the azimuthal component of momentum produced by the fringe field. The reverse process occurs in the exit fringe field, converting from helical motion to simple displacement and divergence. All electrons rotate around their helices at a frequency of 2.800 X 10" revolutions per second, where B is in gauss. For example, an electron entering a 1-kG solenoid field on a parallel path displaced 0.1 cm from the axis is deflected azimuthally through the same angle as if it had traveled 0.05 cm through a 1-kG bend magnet field, about 10.5 mrad at 1 MeV. It has an axial momentum of 2.78 moc and a radial momentum of 0.0292 mOc.It follows a 0. l-cm diameter helical path tangent to the axis, completing a helix period in 0.1 n 2.7810.0292 = 30 cm at 1 MeV. It crosses the axis at 15 cm at an angle of 10.5 mrad. This example is illustrated in Figure 7-8. As the electron accelerates to higher energy, the transverse momentum is retained and the helix stretches out. For example, at 6 MeV the axial momentum is 12.7 mot, corresponding to a 137-cm helix period and axis crossing angle of 2.3 mrad at a 68-cm axial distance from where its trajectory was parallel to the axis at 6 MeV.
:--=@-eAxis of Helix
FIGURE 7-8
. Motion of electron in field of solenoid magnet.
122
CHAPTER 7. BEAM OPTICS OF MAGNET SYSTEMS
A short solenoid can function as a thick lens. For example, a 7.5-cm long I-kg solenoid (with sharply defined pole ends) would cause a ray entering offset by 0.1 cm and parallel to the axis to follow 90" of helix rotation to the solenoid exit. At that point, its radius from the axis would be 0.0707 cm; it would have 0.0206 moc azimuthal momentum and 0.00856 mocradial momentum. In passing through the exit magnetic fringe field, the azimuthal momentum would be canceled. The radial momentum would be retained. The 1-MeV electron would converge toward the axis with an angle of 3.08 mrad, crossing the axis in another 23 cm, a total distance from the 1-kg solenoid entrance of 30.5 cm. At 6 MeV, a 34-cm long solenoid would rotate the parallel entering electron at 0.1 cm from the axis through 90" of the helical path, producing 0.67 mrad radial convergence and axis cross-over at 182 cm from the solenoid entrance. Thus, even a thick solenoidal lens is not very useful in a medical accelerator beyond about the first few million electron volts. Solenoids require much more coil excitation power than quadrupoles for equivalent focusing force, but the transverse excursion of the electron rays of a typical beam is less with a solenoid. Hence, for small aperture accelerator structures, such as high shunt impedance SW medical accelerators, solenoidal focusing is used.
BEAM STEERING COILS A set of four coils is typically located near the entrance of the accelerator structure and another set is near its exit. Each set forms a pair of orthogonal dipole magnets to permit deflecting the beam by a small angle in the x and y transverse directions. The entrance set can correct for error in angle of injection of the beam from the gun, aiming the beam along the axis of the solenoid, which is concentric with the axis of the accelerator guide structure. The exit set can correct for an error in angle between the axis of the solenoid and the proper entrance axis of the collimator into the bend magnet. If the deflection angles produced at the entrance and exit are equal and opposite, the result is a transverse displacement of the beam. The angle and position of the electron beam at the output of the accelerator structure, into the bend magnet system, and at the x-ray target or electron window can be controlled by this steering process, thereby correcting for deviations in the electron beam alignment, such as may be caused by mechanical changes or unwanted magnetic fields. When steering coils are used within a solenoid, they produce a change in the radial component of momentum of the electron, which causes it to follow a larger or smaller diameter helical path in the solenoidal field. This converts to simple displacement and divergence as the beam transits the solenoid exit field. The net steering effect is a function of the number of degrees of helical motion, hence of beam energy. In addition to a set of coils near the entrance and another near the exit of
the solenoid, a single set distributed over the full length of the solenoid may be used, which can have the effect of tilting the magnetic axis of the solenoid. This can be used to correct for misalignment of the accelerator structure axis with respect to the solenoid axis (e.g., due to the winding pattern of the solenoid coils), while maintaining minimum helical excursion, hence maximum clearance from the apertures of the accelerator structure.
BEAM TRANSPORT In order to compute the passage of a bundle of electrons comprising the beam through an array of magnets and drift spaces, a set of simultaneous equations needs to be solved. A program titled TRANSPORT has been developed by Brown- and Careyl2.13 to accomplish these calculations on a computer. In the following sections some conventional terms and formalisms used in beam transport computations are defined. Beam emittance, drift lengths, dipole, quadrupole and sextupole magnets, solenoids and magnetic lenses, and their transport of beams are described. Nonachromatic, single achromatic, double achromatic, and isochronous magnet systems and scanning magnets are then discussed in relation to their use in medical accelerators. Table 7-1 lists definitions of terms commonly used in discussing beam transport magnet systems. The transport of a beam from the object plane to the image plane through a magnet system is analogous to the transmission of light through an optical system. Historically, terms from optics have been adopted for use in the development of beam transport theory. For example, chromaticity refers to color wavelength in optics and to particle momentum (electron energylc) in electron beam transport.
BEAM EMITTANCE Beam ernittance is a convenient way of describing the parameters of a bundle of individual electron rays. The acceptance of a beam transport component or system is simply the maximum beam emittance that it will transmit. These concepts are convenient in understanding the performance of an accelerator or beam transport system and in optimizing their design. The motion of electrons within an rf beam bunch is usually defined relative to the bunch centroid, as discussed, for example, by Banford.4 A set of Cartesian coordinate axes is defined with origin at the bunch centroid, z axis in the direction of motion, x and y axes in the transverse plane, normal to the direction of motion. In bend magnets, the origin follows a central trajectory in the median plane between the two poles of
123
BEAM EMITTANCE
Table 7-1. Beam transport system terminology --
Trajectory Displacement Divergence Dispersion
chromatic^ Magnification Stigmatic
Radial Transverse Focus
The path (orbit) followed by an individual electron ray The distance of an individual trajectory from the central trajectory, measured normal to the central trajectory at any point along it The angle between the slope of an individual trajectory and the slope of the central trajectory at any point along it The spreading of displacement and divergence of individual trajectories solely from their spread in momenta, due to their spread in radii of curvature when bending in a given magnetic field. "Doubly achromatic." Spatial dispersion and its derivative, angular dispersion, of a beam are both zero at a selected distance along the central trajectory from the object plane, independent of beam momentum The ratio of a spatial dimensions of the beam section at the image plane to the corresponding dimension at the object plane "Doubly focused." The image of both the radial and transverse components of the object occur at the same point along the central trajectory. They occur at two different points in an astigmatic image In a plane parallel to the median plane of a bend magnet A surface normal to the median plane and containing the electron trajectory Reconvergence of the object plane beam cross section at the image plane, with magnification
aHistorically, some magnet systems have confusingly been termed achromatic or zero dispersion for cases where the spatial dispersion is zero but its derivative, angular dispersion, is not zero. To eliminate confusion, systems having zero dispersion in both displacement and divergence are sometimes termed "doubly achromatic", permitting the use of the term "singly achromatic" for systems having zero dispersion only in displacement.
the magnet, the x axis is in the median plane and the y axis is normal to the median plane. The state of motion of a particular electron in the beam is then represented by a six-dimensional hypervolume phase space vector
where x, y, z are position vectors and p,, p,,, Ap, are momentum vectors relative to the position and momentum of the centroid electron. Ignoring such effects as space charge forces and gas scattering, Liouville's theorem (see, e.g., Ref. 4) implies that the phase space volume representing the particle states of the beam remains constant. Assuming the components of motion are independent in the x, y, and z directions, the areas of xe,, yp,, zp, planes remain constant. In the transverse dimensions this area is known as the emittance of the beam. It is convenient to assume that this area in phase space is bounded by an ellipse. Each point within the ellipse represents a possible electron ray. Portions of the ellipse area may be void of electron rays but it is assumed that no electron rays have phase space values outside the ellipse. The magnitude of the emittance ellipse is usually defined as the ellipse area divided by T and is usually normalized to nzoc.The area of an ellipse is given by the product of T , the maximum projection onto one axis, and the intercept with the other axis. Thus in Figure 7-9, in the x, p, dimension in phase space, the normalized emittance would be the product of the x-axis inter-
cept xi, and maximum projection onto the p,-axis p,,,,, or maximum projection onto the x-axis x,,, and p,-axis intercept px, divided by mot. For beam transport systems through which the momentum of the beam remains constant, it is conventional to divide px and p, by p, of the centroid electron, giving divergence angles 0 = dxldz and = dyldz. (It is also common to use x' and y' instead of 0 and 0 to designate divergence in phase space.) Thus, in the x,0 dimension in phase space the emittance would be the product of the x-axis intercept and the 0-axis projection of the ellipse containing all electrons in the beam. The term x is usually stated in millimeters and
FIGURE 7-9
. Beam emittance ellipse inx,p,
phase space.
124
CHAPTER 7. BEAM OPTICS OF MAGNET SYSTEMS
FIGURE 7-10 . Current density across the beam in real space for a uniform density in six-dimensional phase space.
0 in milliradians. A statement of normalized emittance would be x0 (pzl~~zoc). Figure 7-10 shows the current density across a beam in real space corresponding to a uniform density in x0 and y@ space, similar to that observed in practice. Figure 7-11 shows two electrons a and b, changing their displacement and divergence in real space and correspondingly moving on a constant area ellipse in phase space. For example, electron a moves from x = 1,0=Oatzl,tox=0.707,0 =0.707atz2tox=0,0 = 1 at z3. Similarly, electron b moves from 0 = 1, x = 0 to 0 = 0, x = 1. In a solenoidal field, an initially upright ellipse would remain upright. In other types of beam transport systems, the shape of the ellipse can change from an oblong tilted one way
I
to circular to oblong tilted the other way, but in the absence of nonlinear forces its area does not change. This is illustrated in Figure 7-12 with limiting ray trajectories and corresponding phase space plots for the x dimension at positions along a field free drift length. The region at minimum beam envelope diameter is termed a waist. It has an upright ellipse in phase space. Figure 7-13 shows phase space plots immediately before and after a thin lens; Om, changes but X,, is constant. Nonlinear forces do not change the area in phase space, but the shape can change to nonelliptical so that a larger effective ellipse is required to contain it. This is illustrated in Figure 7-14. Increasing the energy of an electron by linear acceleration (without transverse forces) does not affect its transverse momentum; hence, its divergence angle decreases but its normalized emittance x0 (p,lm,c) does not change. For example, an electron beam from a 30-kV gun, injected with a maximum beam radius of 0.1 cm at its waist (zero divergence at periphery), and maximum divergence on the axis of 2 100 rnrad would have an axial momentum of 0.348 nzoc, and maximum transverse momentum of 0.349 moc. After linear acceleration to 10 MeV in 100 cm, the axial momentum would be 20.546 moc, and the maximum transverse momentum would still be 0.349 moc, corresponding to a maximum divergence of 0.349120.946 = +1.7 mrad. The phase space ellipse of the output beam would have tilted due to the transverse displacement of the initially axial divergent rays by 0.175 cm over the relativistically foreshortened acceleration length L,:
I
I
-x
a
FIGURE 7-11
I
I
I
I
. Positions in real space and phase space for electrons a and b in a solenoidal field.
Ia
I
I
m a s e space
125
NONACHROMATIC BEND MAGNET SYSTEMS
NONACHROMATIC BEND MAGNET SYSTEMS
FIGURE 7-12
. Tilting of phase ellipse in transiting drift space.
where
L E, E,
= acceleration
length input beam energy = output beam energy =
Isocentric treatment units of early design employed a nominal 90" or 270" nonachromatic beam-bending magnet with the accelerator guide structure mounted approximately horizontally in the gantry. Figure 7-15 illustrates the effect of a simple 90" dipole magnet on the exit beam for entrant beams having an energy spread + AE about E,, or a radial displacement k , or a divergence be. These aberrations, which are easiest to understand individually in the radial plane of simple 90" magnets, can appear in combination and are also present in the transverse plane of magnets. Typically, the beam energy spread is restricted to + 10 percent or less about its central value. The trajectories of the low, central, and high energy components in Figure 7-15a are denoted by I , c, and h, respectively. A single 90" bend magnet, such as shown in Figure 7-15a, is not achromatic. It can bend a monoenergetic beam on axis to a point at the x-ray target but the spread of energies of the actual beam results in a spread of such focal points at the x-ray target as shown. This spreading effect can be minimized by reducing the bending radius of the magnet, restricting the emittance and the energy spread of the beam, stabilizing the operation of components that affect beam energy, or by incorporating a second magnet to provide focusing. However, even with these precautions, changes of energy, as well as variations of the angle and position of the entrant electron beam, will produce detrimental asymmetries in the exit beam and treatment field. As illustrated in Figure 7-16, the principal effect of an energy change of the entrant beam in a nonachromatic 90" magnet system is a lateral change in displacement and angle of the beam at the target with consequent production of asymmetry of the treatment field, especially at high x-ray energies because of the steep slope of the x-ray flattening filter. Energy controlling slits can be located near the output of a 90" magnet, but
I FIGURE 7-13
. Tilting of phase ellipse by thin lens.
Phase Space
126
CHAPTER 7. BEAM OPTICS OF MAGNET SYSTEMS
FIGURE 7-14 . Area of phase space ellipse remains constant from a to b under action of linear forces (Liouville's theorem). The area in phase space remains constant from a to c under action of nonlinear forces but the contour is distorted, requiring an enlarged ellipse to encompass it, hence enlarged effective emittance (from Ref. 4).
Magnet Pole
/
AX r
Electron Beam C
C
r
t
L
*
.-
C
AX
7 l EO- A E
FIGURE 7-15
lf
7
c h Eo Eo+A E
c
. Effect of 90" dipole magnet on exit beam having (a) energy spread, (b) radial displacement, and (c) radial divergence.
C
127
NONACHROMATIC BEND MAGNET SYSTEMS
Electron Beam
max
FIGURE 7-16 . Flattened x-ray field distributions. (a) Symmetrical, electron beam axial at x-ray target, (b) asymmetrical, electron beam tilted at xray target; and (c) asymmetrical, electron beam displaced at x-ray target.
here they tend to become pseudo-targets enlarging the effective focal spot size. The design and performance of the 90" bent-beam Clinac 6 treatment unit have been reported by Austin,] Haimson et al.18 and Horsley et a1.20 and its magnet described by Avery.2 It avoided some of the beam lateral displacement and angle problems at the x-ray target by allowing the lower energy portion of the beam spectrum to sweep by the edge of the target into a shielded cup and by using feedback to keep the high energy portion of the beam spectrum directed at the correct angle onto the x-ray target near its edge. Figure 7-17 illustrates a 270" uniform field magnet wherein a 10 percent energy spread beam comprised of parallel rays is brought to a single focal point in both the radial and transverse planes at the x-ray target, in part by choice of the of the-entrant and exit pole faces' higher energy component h is deflected through a circle of larger radius9 the lower energy component through a circle of smaller radius, but
+
RGURE 7-17 . Spmd in an& of exit bean
by 1700
achromatic Ytriple-focus"bend magnet due to energy spread I to h, of entrant beam.
128
CHAPTER 7. BEAM OPTICS OF MAGNET SYSTEMS
270" Magnet
Magnet Poles
Section dl - d2
FIGURE 7-18
. A 270"magnetic mirror-approxin~ately hyperbolic pole contours.
Magnet sectors
~ a ~ n poles e t
FIGURE 7-19 . A 270' single sector magnet with locally tilted pole gap (gradient shim pole pieces) and adjustable angle input and output pole faces.
ACHROMATIC BEND MAGNET SYSTEMS
both converge on the target at the same point on the central energy trajectory c. However, a change in mean energy of the beam from the accelerator structure will result in a change in mean angle of the beam at the target and, hence, to asymmetry in the x-ray field. This magnet has been termed "triple-focus zero dispersion" because initially parallel cosine-like rays in both radial and transverse planes are brought to a point focus, independent of momentum. It has also been termed "singly achromatic," but it does not comply with the definition of "achromatic" (doubly achromatic) because of the dependence of angular divergence at the image plane on momentum.
ACHROMATIC BEND MAGNET SYSTEMS In order to minimize distortions in flatness of radiation fields due to changes in mean energy of the energy spectrum from linear accelerators, machines that require bending the electron beam employ an achromatic bend magnet system. Taking into account the magnification factor from object to image, the characteristics of the individual rays making up the electron beam are the same after the bend as before it, independent of their momentum spread. With unity magnification this has an effect equivalent to placing the accelerator structure output directly at the x-ray target and in line with the x-ray beam axis. Achromatic magnet systems have been used for decades for translation or deflection of beams from stationary linacs (see e.g., Panofsky et al.,26 Penner,27 and Brown et al.9). In general, more compact systems are required for 360" isocentric accelerators. This need has led to the several different designs of achromatic magnet systems in use in medical electron accelerators. Each of these designs is discussed in the following sections.
SYMMETRICAL 270" SINGLE SECTOR HYPERBOLIC POLE GAP Figure 7-18 shows the earliest concept of an achromatic bend magnet that was sufficiently compact to be applicable to high energy medical electron accelerators. It was proposed by Engels-17and is called a magnetic mirror because particles that traverse it appear to be reflected from the plane at its entry. It is employed in the Brown-Boveri Dynaray treatment units, which are described by Sutherland.30 Its magnetic field increases approximately in proportion to distance in from the entrance plane, the field exponent, n having a value of about - 1 for the first 1 cm in from the entrance plane and -0.8 for the remaining several centimeters. That is, By = Gr-n, where G is the field gradient and r is the distance in from the magnet effective entrance face (with field clamp). Higher momentum electrons travel through a higher magnetic field, lower momentum electrons travel through a lower magnetic field. All electrons entering on the 0" axis coalesce on the 270" axis regardless of their momentum difference. There is no
129
dispersion and the deflection is achromatic. This magnet is capable of faithfully focusing a wide range of momenta as limited by slits S1 and S2 placed at the symmetry plane dl-d2. However, highly precise manufacturing tolerances are required to shape the contoured pole faces. To obtain greater freedom of design and control, locally wedged and stepped pole gap magnets have been developed.
SYMMETRICAL 270" SINGLE SECTOR LOCALLY TILTED POLE GAP Figure 7-19 shows one way to achieve the achromatic beam transport of the "mirror" magnet, by combining uniform and nonuniform field regions in the same magnet sector, as proposed by Enge.16 It focuses a range of entrant momenta, and an input configuration of lateral displacements and angular divergences of the electron beam, to an optically similar configuration at the output focal plane. Its entrant and exit pole face angles a l and a2can be adjusted for optimal radial and transverse focusing of each monoenergetic bundle of rays at the dl-d2 plane. In addition, two adjustable pole sections of the magnet, shown in section view dl-d2, provide an adjustable radial field gradient with 12 < 0. Higher momenta rays travel through a higher magnet field, lower energy rays through a lower magnetic field in this localized 1.1 < 0 region. This radial gradient is controlled by adjusting angle a3to reconverge the different energy bundles or rays into a single spot at the x-ray target. The distribution of rays in this spot are the same as in the beam cross section that enters the magnet. Petersilka et a1.28 describes the employment of this type of bending magnet in the Siemens Mevatron.
SYMMETRICAL 270" SINGLE SECTOR STEPPED POLE GAP Figures 7-20 and 7-21 show another way to achieve the achromatic beam transport of the Enge "mirror" magnet. Brown et al.1 and Tronc31 describe the beam trajectory in this type magnet. In the specific example depicted here, the tnjectory bends approximately 49" on a large radius, pol in a uniform low magnetic field wide pole gap region; approximately 172" on a short radius po2 in a uniform high magnetic field narrow pole gap region; and again 49" on a large radius pol in a uniform low magnetic field wide pole gap region. High energy rays follow a longerpath length than low energy rays, in order to experience the same total bend angle. Rays of different momenta are spread apart at the symmetry plane, where energy selection slits are located. Both cosine-like and sine-like rays have crossovers in the median plane at this symmetry plane, the sine-like loop appearing to tilt with the entrance divergence angle. In the transverse plane, the cosine-like rays have a cross-over at the symmetry plane so the transverse dimension of the beam at the symmetry plane is defined by the sine-like rays. The magnet entrancelexit pole face is at 4.5" to the beam, as in the Enge mirror magnet. The position of the step and the ratio of high/low magnetic fields (as aided by trim coils) are chosen so that the
CHAPTER 7. BEAM OPTICS OF MAGNET SYSTEMS
130
'
Entrance Collimator
Target Plane
'
/ Median Plane
FIGURE 7-20 . A270° single sector stepped pole gap magnet: Cosine-like Cr,sine-like Sx, and dispersion Dx trajectories in median plane and sinelike Sg trajectory in transverse plane (from Ref. 11).
object and image planes are located at a distance from the entrancelexit pole face, providing room for a magnetic field clamp to shape the fringe field and for the entrance beam collimator, exit beam window, and x-ray target.
2700THREE POLE GAP, TWO Cx CROSS-OVERS The illustrations comprising Figure 7-22 a-d, describe the nominal 270" magnet system proposed by Brownlo and used in the Varian Clinac 18 family of treatment units. Figure 7-22a is
a cross-sectional view of this magnet in the median bending plane of the central orbit. It incorporates three uniform field dipole sectors, M,, M,,and M3 with short drift tubes connecting them. A magnetic shunt between poles (not shown) provides a region relatively free of magnetic field for passage of the varticles between sectors. The emittance of the beam entering the magnet system is typical of the output beam of an electron linac in terms of cross-sectional area, divergence, and energy spread. The performance of the system is analyzed with respect to the particle that enters along the central axis with reference momentum po and whose central orbit reference trajectory is shown as a heavy dashed line in Figure 7-22a and b.
ACHROMATIC BEND MAGNET SYSTEMS
Entrance
131
all particles outside a selected momentum band tAp centered aboutpo. An end view of this magnet system in the transverse plane is shown in Figure 7-22c. The entering angular divergence S, and lateral displacement Cx in the beam cross section from the accelerator are reproduced at the target plane (see Fig. 7-22a) with no significant increase in their magnitude. Figure 7-22d illustrates focusing properties in the transverse plane due to the fringing fields of the shaped pole faces along and near the reference trajectory and depict angular divergence Sy and lateral displacement C,. Again, the entering transverse divergences and displacements are reproduced at the target with no significant increase in their magnitude. The dimensions of the entering radial and transverse divergence and displacement are the same in the radial (bending) plane, so a circular symmetric beam from the accelerator structure is reproduced in all its aspects to first order by the magnet system, at its exit, independent of energy. The system is achromatic since both the spatial dispersion and its derivative are zero at the output plane. The energy slits are sources of leakage radiation from stopped electrons. However, this bremsstrahlung is directed away from the isocenter and not directly contributory to patient exposure. A magnetic analysis of the electron treatment beams from a Clinac 18 by Wessels et al.32 gave energy dispersion values of 50.4 to 0.7 MeV full width at half-maximum (fwhm) over the range 6-18 MeV, respectively. These dispersion values include the effect of scatter from the linac thin window and 1 meter of air. A description of magnetic and threshold techniques for energy calibration of high energy radiations has been given by LanzI.23
FIGURE 7-21 . Cross section of a 270' single sector stepped pole gap magnet structure and of edge field clamp (from Ref. 11).
Brown6 showed that the properties of such a system are completely determined to second order by specifying five representative trajectories or paths relative to the reference trajectory. Spatial departures from the reference trajectory by particles of reference momentum are separated into orthogonal radial and transverse components for rays initially at the axis but with divergent trajectories S, and Sy and for rays initially parallel to the axis but displaced from it on trajectories C, and Cy The momentum trajectories D, dispersed by the magnetic field from the reference trajectory in the median plane are for particles that are initially axial and with momenta differing by + A p from the reference momentum yo. Two of these trajectories (see Fig. 7-22a) depict divergence from the central orbit (ST)and lateral displacement from the central orbit (C,) in the median plane. Figure 7-226 is a simplified view similar to Figure 7-22a depicting trajectory D, of momentum dispersed particles initially on the central axis. The trajectories of all particles through the system are symmetrical about a plane of symmetry located at 135" midway along and normal to the reference trajectory. The energy selection slit is placed at approximately the first C,ycross-over to intercept
SYMMETRICAL 270" THREE SECTOR UNIFORM POLE GAP, ONE Cx CROSS-OVER Achromatic systems with three sectors have been used in the past for beam transport in applications not directly related to radiotherapy. Steffen29 shows such a system with three identical sectors (see his Fig. 3.3a), with two Cx crossovers and with one S, cross-over in the midplane. Livingood25 shows such a system (see his Figs. 11-9 to 11-11) with first and third sectors identical and the central sector angle chosen between 90" and 216", depending on the desired locations of the object and image planes. Leboutet24 describes a 270" three sector system applicable to medical electron accelerators. As with Livingood,25 the first and third sectors are identical and the angle of the central sector is selected to provide a suitable drift space between sectors. Two types of design parameters are described, one having two cosine-like cross-overs in the median plane and the other having one cosine-like cross-over. The entrance and exit pole faces of the central sector can be curved cylindrically to compensate for second-order aberrations. The single cross-over design is used in the Therac 20 accelerator, with three identical 90" sectors (see Fig. 7-23).
132
CHAPTER 7. BEAM OPTICS OFMAGNET SYSTEMS
Magnet Coil
Magnet Sector
Reference Trajectory
'.___..'
Plane of Symmetry
Slit
I Entrant Trajectory ' I
,
n
./
Dx
/
A
c..
!
I
I
Pi/
Energy Selection Slit
', I
Central Orbit
I
Reference Trajectory
I
*P
Magnet Coil \
Plane of Symmetry
Beam Apertures
Central Orbit Reference Trajectory
(4
.
FIGURE 7-22 . A 270" three sector uniform pole gap magnet (baqed on Ref. 10). (a) Cosine-like Cx and sine-like Sx trajectories in median plane (b) Momentum trajectories Dx dispersed to maximum at 135'. Energy slit is at Cx focus at 90'. (c) A transverse cross section. (d) Cosine-like Cy and sinelike S, trajectories in transvefse plane.
ACHROMATIC BEND MAGNET SYSTEMS
i
Material Code Tungsten B Lead BZBi
and the second bend is 77". The exit pole of the first sector and the entrance pole of the second sector are parallel to each other but rotated by angles ---ql and + y 2 of 32" from normal to the central trajectory. A quadrupole doublet is placed between the accelerator structure and the bend magnet system to transform the beam phase space ellipse to match the acceptance phase space of the magnet system. The measured dependence on the energy of the output beam parameters is less than 0.05-mm displacement and less than 1.2-mrad divergence per percent energy change.
1
ASYMMETRIC 112 Y2" THREE SECTOR UNIFORM POLE GAP
Fixed Collimator
Im
Scale 10 cm:
Treatment Field 4 Axis -Y
.
I
Adjustable Collimator, Set For 10 cm Field lsocenter Patient Plane
FIGURE 7-23 . Cross section of 270" three sector uniform pole gap magnet showing electron beam sweeping magnet and radiation head. Magnet employs single cosine-like Cx cross-over (from Ref. 3).
ASYMMETRIC 270" TWO SECTOR UNIFORM POLE GAP Figure 7-24 shows a system described by Hutcheon et al.",22 which is designed to minimize the height of the 270" trajectory loop in order to limit the required isocenter height of an isocentric medical electron accelerator. The first bend is 193"
nn,
1I
Target,
;
133
FIGURE 7-24 . A symmetric 270' two-sector uniform pole gap magnet en~ploying193' first bend and 77" second bend. Solid line is nominal energy central ray. Dashed line is higher energy ray (from Ref. 22).
Figure 7-25 shows a system described by Bates? which is used in the SL25 accelerator. This machine employs a drum type gantry, a design that provides more space horizontally than in stand-mounted gantries. The accelerator guide can extend back through the gantry support bearings nearly to the treatment room wall instead of stopping near a gantry stand. There is room to distribute the magnet sectors along the gantry axis direction and still use a relatively long (2.5 m) accelerator structure. This permits using a total bend nearer to 90" instead of the conventional 270°, still achieving approximate achromaticity. The goal is to minimize the height of the trajectory above the x-ray target in order to minimize isocenter height. The system employs identical +45" first and -45" second sectors and a 112 IhOthird sector. The accelerator guide and away from the gantry axis. There is a C, input beam aim 22 lhO cross-over at the plane 4between the second and third sectors. As shown, a second C, cross-over occurs at the image plane F independent of momentum, producing a reduction in beam cross section at the target for initially parallel rays but an increase in angular spread at the target for initially divergent rays. Maximum dispersion of momenta occurs at plane L2, but the energy slit S is located at the entrance to the second sector to minimize x-ray shielding above the third sector, again to minimize isocenter height. The entrance and exit pole faces of the three sectors are rotated by a few degrees to provide a C, cross-over at the object plane. Under these conditions the beam cross-section at the object plane is determined by the input beam divergence S.r,S,., which can be small, especially at high energy. The sector faces are curved to correct for secondorder aberrations.
SYMMETRICAL 180" FOUR SECTOR UNIFORM POLE GAP-ISOCHRONOUS Figure 7-26 shows a system described by Heighway,lg which is used to reflect the beam 180" for a second pass through the accelerator structure of the Therac 25 accelerator. An isochronous design is employed to avoid spreading of the rf electron bunch length from input to output of the magnet. This is
134
CHAPTER 7. BEAM OPTICS OF MAGNET SYSTEMS
FIGURE 7-25 . A symmetrical 112 YzOthree-sector uniform pole gap "slalom" magnet (from Ref. 5) .(a)Cosine-like Cx ray in median plane at nominal, higher, and lower energies. (6)Cosine-like Cx rays in transverse plane.
important to achieve adequate beam current through the energy slits to achieve an adequate dose rate in the x-ray mode, especially at low x-ray energy. The central trajectory bends approximately -6O0, +115", +35" to a symmetry plane and +35", + 115", -60" to return to the accelerator structure axis. To first order, all electrons take the same time to traverse the magnet system, independent of their input displacement, divergence or momentum relative to the centroid electron ray and each electron ray has the same displacement and divergence on exit from the magnet system as on entrance to it. The magnet location can be adjusted to position the returning electron bunch at the desired phase on the accelerating wave. This method of shifting the phase of the return electron bunch can be used to control output beam energy, provided that the bunch length after this bend is short to avoid excessive energy spread. Electrons of different momenta pass through symmetrically located cross-over points before and after the symmetry plane and cross the symmetry plane normal to it. A higher energy ray follows a longer path than the central trajectory before the first and after the second momentum cross-over and a shorter path between them, such that rays of different momenta have the same total transit time, and, hence are isochronous. Figure 7-26b shows the beam envelope in the median plane and Figure 7-26c in the transverse plane, with a waist at the symmetry plane. The relatively large ratio of
magnet gap / orbit length within each sector implies the need for field clamps, with precise adjustment of their positions to tune the focusing properties of the system.
REFERENCES 1. Austin NA: Electron weapon against cancer. Electronics April 6:88-92, 1964. 2. Avery RT: Electron accelerator with specific deflecting magnet structure and x-ray target. U.S. Patent 3,360,647, issued 1967. 3. Bading JR, L Zeitz, JS Laughlin: Phosphorus activation neutron dosimetry. Med Phys. 837-843, 1982. 4. Banford AP: The transport of charged particle beams. London, E. & F. N. Spon, Ltd., 1966. 5. Bates T: Deflection system for charged particle beam. U.S. Patent 4,409,486, issued 1983. 6. Brown KL: A first- and second-order matrix theory for the design of beam transport systems and charged particle spectrometers. Stanford Linear Accelerator Center Report. SLAC 75 (Rev. 3) UC-28, 1972. 7. Brown KL.: Beam envelope matching for beam guidance systems. Stanford Linear Accelerator Center Report. SLAC; PUB 2370: 1-32,1980.
REFERENCES
135
.
FIGURE 7-26 Isochronous 180° four-sector uniform pole gap magnet (from Ref. 19). (a) Cosine-like Crray in median plane and central ray at nominal, higher and lower energy. (b) Relative beam sue in median plane versus path length. (c) Relative beam sue in transverse plane versus path length.
8. Brown KL, DC Carey, Ch. Iselin, F Rothacker: Transport-A computer program for designing charged particle beam transport systems. CERN 80-04 CERN - Service d'information scientifique -RD/437 -2500, Mars 1-251,1980. 9. Brown KL, WKH Panofsky, JF Streib: Method of focusing charged particles to provide zero momentum dispersion. U.S. Patent 3,138,706, issued 1964. 10. Brown KL, WG Turnbull: Achromatic magnetic beam deflection system. U.S. Patent 3,867,635, issued 1975. 11. Brown KL, WG Turnbull, F T Jones: Stepped gap achromatic bending magnet. U.S. Patent 4,425,506, issued 1984. 12. Carey DC: TURTLE (Trace unlimited rays through lumped elements). Fermilab Report No. NAL-64. 1971. 13. Carey DC: New features in TRANSPORT. Fermilab TM1046, 2041.000. Batavia, Illinois, Fermilab, September 7, 1981. 14. Chodorow M, EL Ginzton, WW Hansen, RL Kyhl, RB Neal, WKH Panofsky: Stanford high-energy linear electron accelerator (Mark 111). Rev Sci Instr 26: 134-204, 1955. 15. Enge HA: Effect of extended fringing fields on ion-focusing properties of deflecting magnets. Rev Sci Instr 35:278-287, 1964. 16. Enge HA: Deflecting magnets, in A Septier (Ed): Focusing of Charged Particles. New York, Academic Press, 1967; vol 2, 203-265. 17. Enge HA: Particle accelerator provided with an adjustable 270'
18. 19. 20.
21.
22.
23.
24. 25. 26.
non-dispersive magnetic charged-particle beam bender. U.S. Patent 3,379,91l , issued 1968. Haimson J, CJ Karzmark: A new design 6 MeV linear accelerator system for supervoltage therapy. Br J Radiol 36:650-659, 1963. Heighway EA: Magnetic beam deflection system. Canadian Patent 993124, issued 1976. Horsley RJ, RH Price, JE Saunders, PW Dingwall: Performance of a 6 MeV Varian linear accelerator. Br J Radiol 41:312-3 16, 1968. HutcheonRM, EA Heighway: Anew compact doubly achromatic asymmetrictwo-magnet beam deflection system. Nucl Instr Meth 187:81-87, 1981. Hutcheon RM, SB Hodge: Design and construction of a novel compact doubly achromatic asymmetric 270" magnet system for a 25 MeV therapy electron accelerator. AECL. Chalk River, Ontario, 7057: 1-100, 1980. Lanzl LH: Magnetic and threshold techniques for energy calibration of high-energy radiations. Ann NY Acad Sci 161:lOl118,1969. Leboutet H: Magnetic deflecting and focusing device for a charged particle beam. U.S. Patent 4,046,728, issued 1977. Livingood JJ: The optics of dipole magnets. New York, Academic Press, 1969. Panofsky WKH, JA McIntyre: Achromatic beam translation systems for use with the linear accelerator. Rev Sci Instr 25287-290, 1954.
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CHAPTER 7. BEAM OPTICS OF MAGNET SYSTEMS
27. Penner S: Calculations of properties of magnetic deflection system. Rev Sci Insrru 32: 150-160, 1961. 28. Petersilka E, WE Schiegl: Mevatron linear accelerator and its position between the cobalt unit and betatron. Electromedica 2-3/75: 99-103, 1975. 29. Steffen KG: High energy beam optics. New York, Wiley Interscience, 1965, 1-21 1. 30. Sutherland WH: Stability of a linear accelerator with "ach-
romatic" electron beam bending. Br J Radio1 49:262-266, 1975. 3 1. Tronc D: Device for the achromatic magnetic deflection of a beam of charged particles and an irradiation apparatus using such a device. U.S. Patent 4,322,622, issued 1982. 32. Wessels BW, BR Paliwal, MJ Parrot, MC Choi: Characterization of Clinac-18 electron beam energy using a magnetic analysis method. Med Phys 6:45-48, 1979.
Treatment Beam Production
Production of both x-ray and electron treatment beams for radiotherapy originates in the radiation head. The radiation head (often called treatment head) is the structure from which the useful treatment beam emerges. The characteristics of x-ray and electron treatment beams are strongly influenced by the design of the radiation head. The characteristics that significantly influence radiation treatment are often assessed from the central axis depth dose and isodose curves of representative x-ray and electron treatment fields (Chap. 2, p 33-37). Standard methods of measuring and stating the characteristics of radiation beams have been established and are described herein as well as in several publications31~8~,83-~5 and in Chap. 2 and Appendix B. Characteristics of specific commercial treatment units are cited for illustration and many others are extensively reviewed in identified references in this chapter. The treatment beam production features of multi-x-ray energy linacs are described in Chap. 11. Much of the information of Chap. 8 applies to the modern machines, such as the Clinac 2100C, listed in Appendix B, an extensive survey of commercial linacs and their features. A representative radiation head design is illustrated in Figure 8-1. The head provides a number of beam-shaping, localizing, and monitoring devices. They include a bending magnet if used, fixed shielding, the x-ray target, flattening filter, and a series of single or dual electron scattering foils, often mounted on a large carousel, and finally large movable collimator jaws. Included also is a field light with a sizable mirror for illumination up to full field size and an optical distance indicator (rangefinder) together with a large diameter, parallel plate, transmission type ionization chamber assembly for monitoring of the full field for control, and interlocking. The ionization chamber and its role in dose monitoring and beam stabilization are described in Chap. 9. An alternate head design is illustrated in Figure 8-2.120 The lower portion of the radiation head containing the collimator jaws and mirror can be rotated over a 180" (or larger) range around the beam axis to accommodate fields angled with respect to the gantry axis. The radiation head rotation and collimator jaw movements are usually motor driven and controlled from a cable-connected hand pendant. This hand pendant also controls the speed and direction of gantry rotation and couch
movements. In addition, it controls the field-defining and rangefinder lights together with room and laser patient positioning lights. One type of portable hand pendant may be disconnected from the cable and used to exercise control via an infrared (IR) communications link. Remote control of collimators from the console facilitates obtaining full-field exposures superposed on a port film taken periodically to verify patient positioning during a course of therapy. The full-field exposures portray additional anatomic detail near the tumor to aid in the verification of treatment beam positioning. The radiation head may incorporate a collision-avoidance safety system to stop gantry and couch motion in order to prevent contact with the couch or the patient. Such a system may be activated by a suitably placed guard ring, touch plate,
X-Ray Target Retractable
\
Bending Magnet Assembly
\
Electron Orbit Primaly Collimator
Dual Ionization
FIGURE 8-1
. Clinac 18 'keatnient head. (Courtesy of Varian)
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CHAPTER 8. TREATMENT BEAM PRODUCTION
Electron Beam
-- yz -
X-Ray Target or Electron Window
Primary Electron Scattering Foils
Carousel With Secondary Electron Scattering Foils and 6 MV Flattening Filter Mylar Mirror
Rotatable Dual-Aperture Collimator 15 MV Flattening Filter
I Monitor Ion Chamber
Shutter Wedge
Independently Moveable Beam Limiting Diaphrams
FIGURE 8-2 cal Systems.)
. Philips S U 5 treatment head. (Courtesy of Philips Medi-
or a capacitive sensor. Such a system may, however, lead the operator to place undue reliance on it. The radiation head of contemporary high energy linacs is crowded with components and subassemblies. A cross section of a radiation head with a bend magnet illustrating limited total space and typical space for individual components is shown in Figure 7-1. Accessibility for service often becomes difficult, but a Clinac 2500 "swing away" collimator design, described by Barnes,l3 may significantly improve access. An earlier Therac 20 head design allows one of the magnet half-yokes and associated shielding to be hinged for improved access.168 The Clinac 1800 and 2100C radiation head can be split into upper and lower halves for servicing. The detailed penetration characteristicsof radiation beams for use in radiotherapy are often described by tables of numerical central axis dose depth data. A standard reference covering a wide range of x-ray and particle beam energies has been published by the British Institute of Radiology31 and a related commentary by La Rivierell2 (see also pp 147-148).
GEOMETRIC RESTRICTIONS OF RADIATION HEAD The provision of large field size capabilities (e.g., for mantle fields or total body irradiation) and for asymmetric fields, particularly when employing dual independent jaws, has resulted in an increased diameter of the lower portion of the radiation head to permit full opening of the enlarged collimator jaws. Care must be taken in design of the radiation head dimensional outline to ensure that the treatment beam can still be oriented optimally without interference with the patient, such as for tangential irradiation for breast treatments. The provision of a high x-ray energy capability has required the incorporation of a large bend magnet system which,
with associated radiation shielding, has resulted in increased diameter and height of the upper input portion of the radiation head. Loss and impingement of the accelerator beam on the input electron collimator, energy slits, and bend magnet vacuum chamber walls causes them to act as intense radiation sources, which require compact high density, high atomic number shielding material around and integral with the bending magnet to reduce leakage radiation to acceptable levels.15 Such spurious sources can be quantitatively investigated by accurately positioning radiographic film around the radiation head in conjunction with a defined exposure.33 The provision of multiple treatment modalities with an associated carousel for the interchange of electron scattering foils and flattening filter(s) in some machines, x-ray target@), and has affected the size of the midportion of the radiation head and, hence, requires a large source-axis-distance (SAD) to ensure adequate clearance between patient and radiation head. (Most treatment units are 100-cm SAD.) At least two linac models employ a slide machanism, rather than a carousel, to change between x-ray and electron modalities.l73.195 In some high energy machines, low density hydrogenous shielding materials, often with boron added, have been incorporated in the head for neutron shielding, further increasing its size. Appropriate choice of beam-line materials and precision beam alignment and focusing, will reduce photo-neutron production, and therefore the need for such shielding. The distance from the x-ray target to the top of the radiation head is restricted because of limitations on convenient isocenter height associated with the need for the radiation head to clear the floor. The incorporation of 270" achromatic bend magnets results in improved beam characteristicsbut may raise the isocenter height 10cm or more compared to earlier 90°bend magnet systems. A low isocenter height facilitates the technologist's work in setting up a patient. This requirement has been a challenge in the design of accelerator structures for in-line 4 and 6-MeV machines and in the design of magnet systems for bent-beam machines. These magnet systems are discussed in more detail in Chapter 7. The distance from the x-ray target to the distal surface of the collimatorjaws must be chosen large enough to accomrnodate the various intervening components and to limit geometric penumbra. At the same time, this distance must be small enough to provide room for the external x-ray wedge filter mount, accessory mount and their insert trays with accessories, an auxillary multileaf collimator (MLC), and still provide adequate clearance from the patient and patient table. Provision for two trays (for shadow blocks, wedge filter, or compensator) severely restricts the distance the radiation head can extend from the x-ray target along the beam axis. The use of compensators tailored for individual patients is becoming increasingly prevalent with the availability of convenient methods for their construction and improved treatment planning. The distance from the last tray to the isocenter must be large enough to provide for convenient access to the patient and for clearance from the patient table, for example, when
139
ANCKLARY COMPONENTS
rotating the gantry, as well as to provide enough distance in design. The desire for larger field sizes, improved beam the air to limit the scatter radiation reaching the patient from characteristics, and convenient, functional accessories has the radiation head. This is most easily achieved with a long led to a number of studies and improvements in radiation SAD, typically 100 cm. Although an 80-cm SAD has become head design.165,170,195 a relatively standard distance for 60C0 radiotherapy units, the high radiation intensity of linacs has permitted common usage of 100-cm SAD while maintaining convenient isocenter height. Many 4- and 6-MeV machines in use are of in-line ANCILLARY COMPONENTS design. Progress in microwave accelerator technology has per&tted obtaining an adequate dose rate at ~ O O - C ~ - S A D Some radiation head components are intimately part of the in such machines in the necessarily short accelerator guides. beam transport system and are treated in Chapter 7. These Here most of the power from the microwave source is used include typical bend magnet systems, magnet core and coils, to produce the required beam energy with little beam power electron beam collimators, and energy slits together with the left for providing a sufficiently high dose rate. However, associated vacuum system. Other radiation head components such accelerators provide an improvement in patient clearance or subsystems are treated in this chapter except the dosemeter over the 80-cm SAD of earlier in-line designs. There are ionization monitor chambers, which are covered in Chap. 9. also many 4- and 6-MeV machines in use that employ a Several accessories, which modify treatment beam charbend magnet and hence, have always permitted 100-cm SAD acteristics or aid patient positioning, may be attached to the design. However, at 8 MeV and above, longer accelerating front of the radiation head. Many of these are described in structures are required, so they are usually incorporated into Chap. 12.Most treatment units are provided with a selection of beam-bent isocentric units. Figure 8-3 is a block diagram of wedge filters that are mounted externally. Any accessory that a typical medical linac illustrating the two basic linac concan modify the treatment beam must be appropriately interfigurations: the bent-beam and the in-line, straight-through, locked. A single, 60" wedge, mounted inside the head and
////////
/
/ /
Electron Gun
I
-
Accelerating Structure
t
t.
II
Bending Magnet (Optional)
I
Waveguide
:I
I-----1 I Treatment Head I
I
I
I
I-
(Straight-thru Design)
--
I
__--I
Target Circulator (Optional)
Primary Collimator Flattening Filter
Waveguide Variable Collimator Pulsed Modulator
Klystron or Magnetron Treatment Head (Bent-beam Design)
tt 4
System
-
Support
FIGURE 8-3 Block diagram of typical medical linac illustrating the two basic radiotherapy linac configurations-the bent-beam and the straightthrough designs. In this bent-beam design, the beam emerging from the accelerating structure is bent approximately 90' by the bending magnet before entering the treatment head. In the straight-through design, the treatment head is placed just beyond and collinear with the accelerating structure and the bending magnet is unnecessary.
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CHAPTER 8. TREATMENT BEAM PRODUCTION
remotely positioned by motor control, may also be used as a universal wedge. Beams with effective wedge angles less than 60" are obtained by delivering an appropriate portion of the dose with the 60" wedged field and the remaining portion with an open (unwedged) field. Measurements of such a universal wedge by Petti and Siddonl67 agree well with the model proposed by the manufacturer. Miller and van de Geijn 136 describe a modification of the wedge interlock circuit so as to accommodate additional large-field wedges in a Clinac 18/20. The advent of computer-controlled independent collimator jaws has led to the development of the dynamic wedge with improved wedged-field characteristics (see Ref. 32a in Chap. 12 and p 206). Stevens et a1 190 describe a satellite digital display that provides a remote indication of gantry, couch, and collimator positions at the console of a Clinac 18 treatment unit. Such a remote display can help minimize patient treatment errors and is provided on at least one linac design.
RADIATION SHIELDING Much of the weight and bulk of the treatment head is associated with radiation shielding and x-ray beam collimation components. This is largely determined by the need to limit radiation leakage around the machine outside the useful treatment beam. In an isocentric machine, this weight must be counterbalanced, adding to the overall machine weight and bulk. Alternatively, the machine may be configured with a primary beamstopper, which may be a fixed or retractable shield to intercept the primary beam emerging from the patient. Supporting these heavy structures entails exacting structural rigidity specifications in order to satisfy the geometric stability requirement of the treatment beams. As discussed in Reference 85, regulations and safety standards have been issued regarding radiation leakage measured at a distance of 1 m from the path of the electron beam as it travels from the gun to the x-ray target; its purpose is to protect personnel outside the treatment room wall. A second criterion relates to leakage in the patient plane outside the maximum useful field; its purpose is to limit total body dose in the patient. A third criterion relates to leakage through the collimator jaws, in the area between the useful beam and the maximum useful beam; its purpose is to limit the dose given to the patient's normal tissue immediately surrounding the treatment volume. Dixonso notes that since 90" and 180" leakage is limited in energy, shielding thicknesses at large angles with respect to the forward direction may be reduced. Devanney48 reviewed leakage radiation requirements to the patient area and outside the patient area of several regulatory and standards setting bodies. Conere et al.42 note that x-ray head leakage may vary with energy in dual energy accelerators. All radiation head components that stop a significant portion of the electron beam involves shielding. These include
discrete sources such as the x-ray target, electron beam collimators, and energy slits, as well as more diffuse sources resulting from electrons lost and stopped along the entire electron beam path. When the electrons are stopped, they generate bremsstrahlung radiation, which is largely forward directed at these megavoltage energies. The discrete sources are more intense and more localized but are usually easier to shield. Locating the shielding as close to the source as feasible will reduce its volume and mass for a given attenuation. The design of nominal 270" bending magnets can affect isocenter height depending on where beam energy defining slits are located.15 An optimal beam focus location from electron optics considerations may place the beamsstrahlung-producing slits at a location where the shielding requirement would increase the radiation head dimension in the direction away from the isocenter and, hence, increase isocenter height in order for the radiation head to clear the floor. High density, high atomic number materials (e.g., lead and tungsten) are commonly used for radiation head shielding. Tungsten and heavy metal alloys are more costly than lead but their higher densities (- 18vs. 11.3 gIcm3) may dictate their use since such shielding can be placed closer to the radiation sources, target, and energy slits, saving space and mass. However, tungsten, a costly brittle metal, is harder to machine when compared to lead, which is easily shaped, typically by casting, but is more easily damaged. An optimal radiation head design incorporates both metals. Depleted uranium (depleted in the fissionable isotope U-238) is an excellent absorber of x rays with a high density of almost 19 g/cm3, and an atomic number of 92 and structural properties equivalent to mild steel.21 Depleted uranium has been used at lower energies but its residual radioactivity, the necessity to electroplate its surface to control scaling, and increased regulatory requirements have combined to raise its cost and restrict its use. Uranium is not used above 6 MeV because of increased photoneutron production. Methods of measurement and characteristics of leakage radiation for a Clinac 18 treatment unit have been described by Lane et al.108 and earlier by Capone and Karzmark33 for 6-MV x rays. Other aspects of head leakage radiation are covered by a number of investigators,42,48,50,78.97,150,178,195,201
Shielding the head for megavoltage x-ray therapy is the dominant need when contrasted to electron therapy. This follows for several reasons. First, the beam current requirement for electron therapy is significantly less, often by a factor of several hundred to as much as 1000. Electrons are less penetrating than x rays and some are absorbed completely in the shielding material by ionization losses. Others are converted to energetic x rays by the bremsstrahlung process, but the process has low efficiency. At x-ray energies of 15 MV and above, an increasingly significant yield ~f'~hotoneutrons is produced by components in the radiation head. Here, low atomic number (Z) materials loaded with boron are effective radiation shields. Borated poly(ethy1ene) has been used as neutron shielding in one head
ANCILLARY COMPONENTS
design. This neutron shielding was incorporated at an earlier date when neutron protection standards were expressed as effective dose in roentgen equivalent man (rem). This was a more stringent requirement for neutrons by a factor of about 10 compared to the present standard expressed as absorbed dose in Gy and presented extreme containment problems above about 25 MeV. The topics of neutron leakage and radioactivation are treated on pages 150-151 and in Chap. 14, page 252.
BEAM COLLIMATORS The treatment head provides two treatment beam collimators for x-rays as shown in Fig 8-1; a fixed primary collimator, and an adjustable secondary collimator, which is also termed the beam limiting device or "jaws." The primary beam collimator defines the maximum angular spread of the x-ray beam. The opening through it is in the shape of a truncated cone or a four-sided pyramid. A conical primary collimator may be fixed in the head but a four-sided pyramidal one must rotate in order to preserve correct orientation of its sides relative to the jaws of the beam limiting device. For the corners of a 40 X 40-cm field at 100 cm SAD not to be restricted, a half-angle of about 16" would be required for a cone or the comers of the pyramid. Typical practice is to provide a 14" half-angle, corresponding to a 5 0 cm diameter at 100-cm SAD, a 35 X 35-cm fully square field, and a 40 X 40-cm maximum field with comers clipped by several centimeters. Tungsten is the material of choice for both the primary and secondary collimators. Space is at a premium in this location and tungsten collimators take up less space than lead. Also, where the x-ray energy is high enough for significant neutron production, tungsten attenuates the neutrons to a much lower energy than does lead. For example, the average energy of photoneutrons arising from 15-MeV incident energy electrons is reduced to about 0.6 MeV in 10 cm of tungsten and only to about 1.5 MeV in 10 cm of lead.131 A dual x-ray energy linac will have at least two different flattening filters in order to provide flat fields at two energies.53J20.176 The simplest flattening filters are shaped somewhat like a volcano, the more pointed cones being associated with higher energies. In some machine designs, they are inserted coaxially into the primary collimator to preserve space. In one dual energy design, two primary collimator filter assemblies are mounted on the carousel and selected by rotation of the carousel.53 The adjustable collimators consist of two pairs ofjaws, one above the other and at right angles. Their defining edges often traverse arcs, or approximate arcs, such that their inner faces are approximately tangential to the radiation beam emanating from the x-ray target thereby reducing penumbra. They define the size of the x-ray trtatment field and act in combination with the electron applicator to produce a properly shaped electron treatment field dose distribution. They are made of lead, tungsten, or depleted uranium so as to limit transmission of the primary x-ray beam to about 1hto 1 percent. Because of scatter
141
from tissue in the useful beam, the dose to the tissue shadowed by the jaws is often significantly higher than would result simply from leakage radiation through the jaws of the beam limiting device. For small fields, most of the shadowed area is covered by two jaw thicknesses, one from each pair, and the primary transmission is very small. For large fields, most of the shadowed area is covered by only one jaw. These and other leakage radiation considerations have been reviewed by Devanney.48 The contamination of electron and x-ray beams within and outside the useful beam is treated on page 150. In normal use, each pair ofjaws is coupled to provide symmetric rectangular fields centered about the axis. However, for some therapy techniques, independent motion of the jaws of one or both pairs is desirable in providing asymmetric rectangular fields. For example, many breast and other therapy techniques half-block one of two pairs of abutting-opposed fields along the beam axis. Positioningone independentjaw along the beamaxis to provide the half-blocked field eliminates beam divergence without the need for a larger, heavier inconvenient external beam half-block. This simplifies the abutment of adjacent fields with negligible field overlap but necessitates precise patient positioning. Such a beam is also useful for shallow tangential arc treatment minimizing the dose to sensitive underlying tissues (e.g., lung). Independent jaw motion is accomplished by uncoupling the symmetrical motion of a pair ofjaws and independently positioning eachjaw with one jaw edge left on axis forthe half-blocked field. Both jaws may also be independently positioned, allowing one jaw of the pair to cross the central axis, to define fields where both field edges are on one side of the central axis. Dual independent collimators, allowing independent motion of all four jaws, provide additional flexibility in blocking offset field edges. However, this necessitates very large jaws and a large secondary collimator diameter in order to shield all areas subtended by the primary collimator. The applications and physics measurements for the independent jaw system of the Clinac 2500 have been described by Loshek.ll9The dosimetry, treatment planning, and specification of these asymmetric fields are more complex than for the symmetric fields case.39.l~,lo2.159.160Khan et al.lmnote that the effect of the independentjaws is similar to that of secondary blocking. They have modified their computer program for generating isodose curves for Clinac 2500 asymmetric fields. Palta et al.160 point out several unusual dose distributions available via asymmetric arc rotation. A simplified technique for treating breast cancer uses both pairs of asymmetric jaws and a single set-up point without the need for couch angulation.180a Multileaf x-ray collimators have been developed to closely encompass irregularly shaped tumors by an alternative method than by constructing individual Cerrobend shields. In one example, 64 independent leaves comprise one pair of jaws. They can be rapidly set under computer control and can be programmed to follow the changing tumor outline for dynamic therapy. Such a multileaf collimator permits leaf spacing equivalent of about 1 cm at 100-cm TSD. They are discussed in more detail in Chap. 2, pages 41-43.
142
CHAPTER 8. TREATMENT BEAM PRODUCTION
FIELD LIGHT AND RANGEFINDER Patient treatment fields are positioned with the aid of reference points or lines. These are often marked indelibly on the skin with the patient in the treatment position. The center of the field may be marked with a cross, the edges with dashes or lines, all of which aid the set up procedure. These markings are affixed during simulation once the field sizes, field positions, and angulation are agreed upon. They are referenced to the internal anatomy revealed in the localizing radiographs taken with the simulator and to the bony landmarks externally identifiable. The markings remain on the patient's skin through the course of treatment but may be removed and redrawn as appropriate. For example, a tumor may shrink in size during a course of therapy and then be treated with a reduced field size. The field defining light and rangefinder (see Fig 8-1) provide convenient visual methods for correctly positioning the patient for treatment using the reference markings. The field light illuminates an area that coincides with the radiation treatment field on the patient's skin. The field size is defined by the 50 percent x-ray isodose contour extended perpendicular to the incident normal surface from the depth of dose maximum. The spatial coincidence of the x-ray and light fields is typically within 2 mm. The center of the field is marked by the shadow of a small cross located on the beam axis near the light source or on the Mylar window covering the collimator opening and illuminated by the field light. A high intensity lamp with the filament approximating a point source is needed to provide bright, sharp images of the field with minimal penumbra. This lamp is located off the beam axis, and its light is reflected by a mirror placed on the axis as shown in Figure 8-1. The geometry is such that the filament appears to be located at the effective x-ray target focal source. The lamp involves conflicting requirements: high intensity and a small filament size, as well as long bulb life and mechanical stability of the filament. These requirements may be best satisfied by some slide projector lamps, which have a small filament, a quartz envelope, and are halogen filled. In general, a larger source results in a brighter field and a larger penumbra; a smaller source results in a dimmer field and a smaller penumbra. A thin metal or front-surfaced glass mirror fixed in position on the beam axis is often used as the field light reflector if only x-rays are employed. However, if electrons are available, it would unduly scatter the electron beam and must be moved out of the way for electron therapy. The metal or glass type mirror may be translated out of its normal position or located on a carousel, which is rotated to remove the mirror from the field. In either case, such a mirror must be repositioned with spatial precision in order that the coincidence of the light and radiation field is preserved. Alternatively, the mirror can be in the form of a thin silvered or aluminizedplastic film, which only minimally scatters the electron beam and, hence, can remain fixed in position. The field defining lamp assembly must incorporate convenient, 3-D adjustments for positioning the filament at the effective x-ray
source location. Some field light designs incorporate fiber optics with the source size being defined by the output size of the fiber optic, not the bulb filament. Similarly, the spot location is that of the output fiber optic, not the lamp filament. The quality assurance (QA) program will usually include weekly assessment of the congruence of x-ray and light fields. Often, the exit opening of the radiation head may be covered with a thin transparent film or other plastic lamina to provide a fiducial reference surface and to protect the interior of the radiation head from dust and falling objects during undercouch treatments. The retrospective installation of such a cover results in small changes in x-ray beam characteristics. Compton electrons emitted from the film increase the skin dose and modify the depth dose distribution. The film can result in significant changes in electron output, which varies with both energy and field size.4 The beam central axis reference (cross hair) can also be built into the field light projector or can be crossed wires in the light beam. The rangefinder is used to place the patient at the correct treatment distance. It may be located on the outer front or back edge of the treatment head or on the gantry. It projects a centimeter scale whose image on the patient indicates the target skin distance (TSD). The image of the scale intersects the beam axis at the isocenter and indicates the normal treatment distance (see Fig. 8-1). If the skin is located nearer the target, as in isocentric techniques, a smaller value of TSD than normal is imaged and displaced from the axis. Larger than normal values of TSD will be displaced from the axis in the opposite direction. The rangefinder is also operable for electron therapy. A slot is provided in some electron applicators to accommodate the projected rangefinder image.
ELECTRON THERAPY Since the requisite beam current for electron therapy is typically many hundred times less than for x-ray therapy, most shielding problems, including those for neutron shielding, are comparatively insignificant even at the higher energies. One is concerned clinically with providing: 1. Wide, flat electron fields with modest penumbra. 2. Relatively low surface dose. 3. Deep 80 percent depth dose for each selected energy. 4. Rapid fall-off of dose with depth on the distal side of the depth dose curve. 5. Small amount of contaminating x-ray radiation. 6. Negligible applicator leakage. Achieving and assessing these electron beam characteristics have been the subject of a large number of studies. 4,16,20,29,30,46,56,78,81,88,92,94,99,101,114,116,126,127,134,135,137,145149,156, 160a172.177a,180,183,184,185,187,192,195.1201
Specific commercial
ELECTRON THERAPY
equipments may often be identified from these generic references by manufacturer, model, or energy. In addition, specific more specialized topics are discussed and referenced under individual chapter sections (e.g., electron energy measurements, spectra, dual x-ray energy linacs, and independent collimator jaws). Performance specifications for electron and x-ray beams for a representative treatment unit, the Clinac 18, are given in Table 2-1. Figure 8-4 shows an example of the basic beam subsystem for electron therapy. The x-ray target has been retracted and the carousel rotated to the appropriate scattering foil position for the electron energy to be used. The ion chamber monitors this scattered electron beam. The accessory mount is attached to the radiation head together with the electron applicator for the field size chosen into which a shaped cut-out is inserted. The dosimetric characteristics of electron applicators have been studied by a number of investigators.38.94.97.98.116,175,183.192,205They have been greatly improved from earlier implementations in regards to the factors listed above.
Electron Beam
:
T I I-
X-Ray Target
Primary Collimator
Scattering Foil \
Flattening Filter
143
Bell and Waggenerls describe a method for rapid determination of the energy of electron treatment beams from medical linacs. Anderson and George839 suggest some modifications of the early, range-energy equation developed by Markus128 for scanned electron beam energy calibrations. The latter was derived largely from betatron data for scattered electron beam energies up to about 15 MeV. Johnsen 90 describes how wedge-shaped absorbers can be used with ion chambers to measure electron energy. Almeida and Almond47 describe how Cerenkov radiation from nuclear reactions can be used to determine electron beam energy. Film analysis of electron depth dose has also been employed for energy determinations.57 Dosimetric measurements for scanned and pulsed linac electron beams are more difficult than for scattered beams because of the moving spatial, time sequence of pulses and more difficult chamber saturation.41 Ertan et a1.52 describe a scan-triggered, computer-controlled measurement system. Nafstadius et a1.142 describe dosimetry measurements for a computer-controlled, scanned-beam and an associated data acquisition system. This equipment has both electron and x-ray beams with energies up to 50 MeV from a medical microtron. In a later theoretical study, Huizenga and Storchi79 indicate how measurements of penumbra widths of scanned electron beams in air can be used to generate electron distributions at the surface of the patient. These, in turn, serve as input for calculation of patient dose distributions for radiation therapy treatment planning. The clinically relevant electron energy is often specified by the depth of the 80 percent relative ionization on the central axis on the distal slope of the depth dose curve (see Fig. 2-4). Central axis depth dose measurements have been carried out on specific accelerators noted in the preceding references and by the Radiological Physics Center on over 70 electron-producing machines used in radiation therapy.101 These latter data appear consistent for each machine model and nominal energy. However, the data show that depth dose distributions can vary significantly among different machine models for electron beams having the same nominal energy. The shape of electron depth dose curves can vary widely at shallower depths for the same nominal energy, particularly from contaminants in the build-up regions. Characteristics of electron beams from a medical microtron have been described by George et a1.61 and others.14,25-27 Those for a scanned-beam 25-MeV linac are described by O'Brien et a1.154 and different scanned beams by OtheTS.34,52,168,181,188
Patient
FIGURE 8-4 Beam subsystem for electron beam therapy. Cross section view including central axis of the beam.
Usually, electron applicators are attached to the treatment head with the x-ray collimator jaws set and interlocked to subtend fields a few centimeters larger than the field defined by the applicator. There is accompanying improvement in electron field flatness from this procedure, preferentially at shallower depths. An excessively wide setting may create radiation leakage problems outside the treatment volume, as in the case of pregnant patients for one particular linac,zos as well as more generally in early
[email protected] Schniederl83 evaluated such leakage radiation for another model of linac, and Jones94 suggests that the leakage may be a specific
144
CHAPTER 8. TREATMENT BEAM PRODUCTION
machine's characteristic, which should be evaluated by individual users. Often, a lead or Lipowitz's alloy78 cutout defines the final electron field size and shape. It is placed in the end of the applicator at or near (e.g., 5 cm) the patient skin surface. The dosimetry of such shaped small and irregular fields has been described for electrons from 4 to 10 MeV.145 The systematic variation of small field, electron beam characteristics, particularly the output factor, with standard cone sizes and Cerrobend cut-outs, has been studied extensively at 6, 10, and 20 MeV.177a Electrons scattered off the applicator walls improve flatness at the periphery of the field at a shallow depth, but with less penetrating electrons. However, such fields have poorer flatness at greater depth. A continuously variable segmented electron beam collimator has been described by Robinson and McDougall,l80 and at least one manufacturer makes such an option available. An alternate approach is to provide rectangular electron field definition by attaching four lightweight, quick-connect trimmers to the x-ray collimator jaws. One electron beam collimating, flattening, and monitoring system for a 35-MeV betatron has been described by Svensson,l92 and another by Lindskoug and Dahler.116 Multileaf collimators for electrons are described in Chapter 2, pages 41-43. The provision of electron treatment fields defined in size by electron applicators or collimators is provided by scanning a pencil beam of electrons over the field area or by scattering the static axial electron beam with one or more thin foils as described in the following sections.
ELECTRON SCATTERING SYSTEM The electron beam leaving the bend magnet (or accelerator structure for straight-ahead machines) is about 3 mm in diameter (Fig. 8-3). Scattering foils (or a scanning system) are used to spread this beam to the much larger area required for therapy. Most studies find that a single, high atomic number scatterer will adequately flatten small fields up to about 10 cm in diameter for low electron energies up to about 10 MeV. Additional steps are taken at higher energies and for larger fields. Providing several different scatter thicknesses, often in a carousel, facilitates optimization of beam characteristics for different energies.127 A dual-foil scattering system, with a few centimeters or more between the two foils, significantly improves electron beam flatness characteristics with reduced x-ray contamination, particularly above 15 MeV and for fields 15 cm in diameter and larger.195 Such a dual-foil scattering system is illustrated in Figure 8-5. The first high atomic number (Z)foil is selected to minimize energy loss for a given scattering distribution and the second scatterer made of a low Zcomposite, thicker on axis, functions more as a field flattening filter preferentially scattering electrons to the periphery.195 The thicker portion of the second scatterer may be in the form of a high Z "button" on a low Z foil. Hence, the electron applicator in such systems primarily serves to define the field size and only
Narrow Beam
11 9
8
First Scatterer
Second Scatterer
I
I
Both Scatterers One Scatterer Only
FIGURE 8-5 1977).
. Double scattering foil system (adapted from Brahme
affects flatness secondarily. Athorough analysis of the dual-foil scattering system has been provided by Mandour and Harder.126 The interdependence of electron dose rate, energy, field size, effective source size, distance, and beam scattering or scanning, is more complex than for x rays and should be measured for the individual treatment unit. A large effective electron source size will produce a much larger amount of wide-angle, outside of collimator-limited electrons. It and other parameters for electron scattering have been the subject of a number of s&dies.35.87,101.1~aa197 The location of the effective electron source is usually different from the effective x-ray source location. The effective position of the source for electron therapy is less well defined than for x rays and has been studied by several investigators.35+87,97,197Where scatterers are employed, it has been useful to define a virtual point source location, which differs from that of the scatterer. Its location is dependent on machine design and varies with SSD, field size, and electron beam energy.
ELECTRON SCANNING SYSTEM Most medical electron accelerators employ a pair of metallic foils to scatter the bend magnet output electron beam for electron therapy. Brahme27.28 describes a technique in which a first high Z foil produces a Gaussian distribution of scattered electrons and a second small diameter foil located a few centimeters from the first foil rescatters the central portion of the Gaussian distribution to flatten the electron dose distribution at the normal treatment distance. This technique is satisfactory for electron energies up to about 25 MeV at field sizes up to about 25 X 25 cm at 100-cm SAD, producing x-ray contamination
145
ELECTRON SCATTERING SYSTEM
of less than 5 percent and falloff of depth dose from 80 to 20 percent quite similar to the falloff depth dose of scanned beams. The required thickness of high atomic number scatterer increases as the square of energy and as the square of maximum field size. Thus, electron scanning is justified for large fields at energies higher than 25 MeV to avoid deterioration of the shape of the electron depth dose curve and to avoid excessive x-ray contamination. Historically, mechanically translated and rotated magnets were first used for scanning the electron beam in an isocentric radiation head, as described by Rozenfeld et a1.181 for a 70-MeV stationary linac. Rosenfeld et a]., developed a scanned pencil beam system for arbitrarily shaped fields defined by a full size template, in lieu of using scattering foils.34.188 This early isocentrically directed and scanned beam, up to 50 MeV in energy, employs a large, complex nonachromatic beam transport system with attendant treatment beam stability problems. It combines a linear translation of the magnet up to 21 cm along the direction of the gantry axis with an indexed rotation of the gantry to provide 5-mm spacing between scan lines at the skin surface. A treatment is completed in one scan series covering the field, and different electron energies can be used in different portions of the field. Electronic scanning of the electron beam in an isocentric machine was first used in the Sagittaire isocentric radiation head, which was fed from a stationary linac at energies up to 32 MeV in standard machines and up to 40 MeV as an option. Briot et a1,29.30 investigated the scanned electron beam of a Sagittaire 35-MeV treatment unit. They concluded that an adjustable outboard metallic collimator, which could be attached to the x-ray jaws, improved the depth dose and dose gradient at the edges of the field. In regard to this machine, Leboutetlls states that to achieve wide fields at these energies, scanning of the beam had to be used, in addition to the scattering by the window, radiation monitors, and thin scattering foils. Because of the pulsed nature of the beam, the scanning produced a sequence of spots and some foil scattering was found necessary to provide sufficient spot overlap for adequate uniformity of dose distribution. Having developed electron scanning for the 32-40-MeV machine, the manufacturer continued this technique for the Satume type 20-MeV machine as described by Aucouturier et al.10 and Azam et al." and its sequel, the Therac 20 type 20-MeV machine (see Fig. 7-23). With the introduction of the 50-MeV racetrack microtron for radiotherapy, electron scanning is again used to obtain large fields at such high energies without the excessive deterioration of depth dose falloff and excessive x-ray contamination that would result if thick scattering foils were used.27.28 To achieve such scanning in the Therac 20, saw tooth modulated currents of 0.615 Hz in the x direction and the 4 Hz in the y direction are applied to a four-pole sweeping magnet below the 270" bend magnet, as depicted by Badingl2 and shown in Figure 7-23. The amplitude and polarity are controlled to produce a raster scan in the electron treatment mode. Both the x-ray target and the sweep magnet are in air, just outside the electron beam vacuum window, the x-ray target
being retracted for electron mode. Incidentally, a four-pole magnet may look physically like a quadrupole but it functions like a rotatable dipole, as illustrated in the following pole polarity sequences (see also Fig. 7-6).
N
S
Quadrupole
S
N
N
S
or S
N
Pfalzner and Clark168 carried out a depth-dose study of the scanned electron treatment beams of the Therac 20 and concluded that they did not differ appreciably from those of the foil-scattered electron treatment beams of the MM-22 circular microtron reported by Svensson et a1.193 O'Brien et al.l54,155 similarly studied the scanned electron beam from a Therac 25. Others have compared the Siemens betatron and Sagittaire linac electron beams.46
MICROTRONS VERSUS LINACS FOR ELECTRON THERAPY Microtrons, having a wide energy range, have come to be employed for electron therapy.24,25,27,61A controversial aspect of linac versus microtron treatment beams concerns dependence of the characteristics of the electron beam depth dose on the energy spread of the electron beam. Brahme and Svensson24 contend that the microtron beam has a depth dose significantly improved (sharper build-up and steeper fall-off regions) over that of the linac, which they attribute in part, to the narrower energy spread of the microtron beam. Fregene56 suggests that a small difference in energy spread, which is of the same order of magnitude in the microtron and linac beams, should not lead to appreciable differences in depth dose at 10 MeV. Some of the disagreement between observers might be explained on the basis that measurements were performed on different accelerators having different treatment heads (different collimators, scatterers, geometry, etc.) using different phantom materials and dosimetry techniques. Measurements by Johnsen et al.92 of electron beams from a single linac operated at nominal peak energies of 6 and 12 MeV appear to have resolved this controversy. Based on magnetic analysis of the beams, they compare narrow energy spread beams (0.1-and 0.2-MeV fwhm) with broad energy spread beams (0.8-and 1.2-MeV fwhm) for 6 and 12 MeV, respectively. They conclude that, when consistent measurement techniques are used for the beams tested on the same beam collimator system, the electron depth-dose characteristics are not significantly affected by these relatively large changes in the width of the accelerator's energy spectrum (see also Appendix, Fig. A-7). Earlier, Bjamgard et a1.20 found no significant difference between the Mevatron XI1 electron depth dose curves and those from a microtron.20
146
CHAPTER 8. TREATMENT BEAM PRODUCTION
X-RAY THERAPY Megavoltage x-ray characteristics and the equipment that provide them constitute a widely studied area of radiological sCienCe~3,5,32,33,36,37,4O,42,49,584O,68,72,73,75,77,8O,8293,104,122,123,13O,
138,146,152,158,1M),160b,163,171,174,177.182,185,196,202 Specific c o m e r -
cia1 equipments may often be identified from these generic references by manufacture, model, or beam energy. Most of the high or dual x-ray energy equipments also provide electron therapy. Figure 8-6 shows an example of the basic beam subsystem for x-ray therapy. Typically, the x-ray field size is continuously variable from zero, or within a few centimeters of zero, usually to a maximum of 40 X 40 cm at 100-cm TSD. The bremsstrahlung beam from the x-ray target is limited in maximum field size by the primary collimator. The comers of the largest field sizes may be cut off by this restriction. The carousel is rotated to bring the appropriate x-ray flattening filter into position. The ion chamber monitors the flattened x-ray field. The secondary collimator jaws are set to the correct field size, and any necessary wedge, block, or compensator is mounted in the appropri-
4'
Electron Beam
Primary Collimator
Forward Peaked X- Ray Beam
Flattening Filter Scattering Foil
Carrousel Ion Chamber
Secondary Collimator
-
~-4
1 ;
;; ; ;
; !
I
/
Flattened X-Ray Beam -
-I I I I I I I 1 1 '
I
I I 1
I . I I
I. I
I
'
I \-I I \ 1r--i---i--,---l I I I ) I I I I I I 1 1 I
I I I I I I 1
I I I I I I 1 1
I I I I I I 1 1
Slot For Wedges, Blocks, Compensators
I l I l l \ 1 1
~
Patient
.
FIGURE 8-6 Beam subsystem for x-ray beam therapy. Cross section view including central axis of the beam.
ate accessory mounting slot. The x-ray energy of a treatment unit is often expressed and compared with other units by specifying the percent of central axis depth dose for a 10 X 10 cm field at 100-cm SSD at 10cm depth in water. Representative values are 67 percent at 6 MV, 70 percent at 8 MV, 73 percent at 10 MV, 77 percent at 15 MV, and 80 percent at 18 MV. The penetrability of an x-ray beam may be defined by the ratio of ionization measurements made with a fixed sourcedetector distance at depths of 20 and 10 cm of water.' Performance specifications for x-ray beams for a representative treatment unit, the Clinac 18, are given in Table 2- 1. Acomparison of the x-ray central axis depth dose data with published data for a number of commercial linacs has been presented by Gastorf et al.58 and others.31.36.171 Among their individual conclusions, they recommend that physicists verify the applicability of published data to their machine rather than rely upon a set of data measured independently and assumed to be valid. Characteristics of the photon and electron beams produced by dual energy linear accelerators have been reported by a number of inve~tigators.6~13~53.82.93~1~~1~a+16~~176 A more general reference on the parameters of electron and photon treatment beams is Supplement 17 of the British Institute of .. Radiology (BIR)31 and a related commentary of La Riviere.112 The latter questions the Supplement 17 method of correlating the penetrative quality of high energy x-rays with depth dose at 10-cm depth by the use of nominal accelerator megavoltage MV. He suggests instead that the dose-weighted average energy of the filtered beam be used as a true measure of x-ray beam quality at 10-cm depth. Where applicable the earlier BIR Supplement 11 provides a more consistant set of data. Megavoltage x-ray beams exhibit an increased surface dose in the build-up region and a shift of Dm,,towards the surface as the field size is increased. Recent investigations indicate that these effects are largely due to low energy electron contaminants originating in components of the treatment head, primarily the flattening filter,19.124 Petti et al.166 identified the sources of electron contamination for a 25-MV photon beam from a Clinac 35. Such contamination can significantly alter the build-up characteristics of x-ray beams. They conclude that, for normal treatment distances, the flattening filter and monitor chamber contribute 70 percent of these electrons. Treatment heads can be equipped with electron filters to attenuate this component.23,74."8 Reductions of 10-20 percent in surface dose and an increase in depth of Dm,,from 2.5 to 4.5 cm at 25 MeV have been observed for such filters.74 These reductions may be negated however, when using shadow trays, wedges, and so on. Moyer has identified systematic patient x-ray dose errors for 4 and 10-MeVbeams associated with elongated rectangular collimator fields.139 This "collimator exchange effect", which depends on which movable collimator pair forms the larger or smaller field dimension, approaches 2.5 percent for highly elongated fields. A smaller dose is delivered when the upper collimator pair adjacent to the monitor chamber forms the smaller field dimension. Hence, the corresponding correction
X-RAY THERAPY
factor may be applied to the dosemeter readings. This effect appears to be associated with secondary scatter from the collimator into the treatment field and not backscatter into the monitor chamber, which contributes only a vanishingly small amount as shown by Huang et a1.77 and by Watts and Ibbott.203 Kubo et al.lo~,lo6found that backscatter into the monitor ionization chamber can vary with radiation head design. It ranges from less than 1 and 2 percent for 6 and-18 MV photon beams, respectively, for a Clinac 1800 to as large as 7.5 percent for a Therac 20. The x-ray beam characteristics of the Therac 20 have been reported by Patterson and Shraggel62; those for a Clinac 18 by Connor et a1.,43 those for the 18-MV Mevatron 77 by Palta et a1.,158 those for the Clinac 1800 by Findley et al.>3 the Therac 25 by Aldrich et a1.> the Phillips SL 25 by Palta et a1,160b and those for the Clinac 2500 by Krithivas and Rao.104
X-RAY TARGET AND FLATTENING FILTER The x-ray target and flattening filter combine to determine important characteristics of the x-ray treatment beam. The bremsstrahlung beam emerging from the target has fluence, energy, and angular distributions. These distributions are modified in important ways by the flattening filter. Figure 8-7 shows the isodose curves for a 20-MV x-ray beam without and with a beam flattening filter. The detailed effects of the filter on beam flatness for several alignment situations are illustrated in Figure 7-16. An angular or a lateral displacement of the beam on the target produces the unflattened distributions shown. Both effects may be present
-
FIGURE 8-7 Isodose curves for 20-MV x-ray beams; (a) without and (b)with a bean1 flattening filter. (Courtesy of W. J. Meredith and J. B. Massey.)
147
at the same time, complicating the analysis. Even the changing earth's magnetic field encountered in gantry rotation may have an effect. A mechanically stable structure together with an automatic beam steering system such as described in Chap. 9 may be used to maintain the beam accurately centered on the axis of the flattening filter. Rigid linac mechanical construction is a distinct advantage. A number of studies have focused on x-ray flattening filters, target selection, bremsstrahlung, and beam penetration (e.g., percent depth dose at 10 cm o r depth of 50 percent depth doSe)~22,26,27.34,36,44,45,54,62,65d7775-78,86,9,95,97,lO7,l13,117,121,129,
130,143,144,152,167,170,171,179,191,195
X-ray spectra have been investigated by several groups3-18-75,13O9138 and are an important parameter in the design of flattening filters and the characteristics of high energy x-ray beams. An early discussion of beam flatness of seven different treatment units including betatrons, 60C0, and linacs was included in a study by Chan et al.36 Kase and Svensson97 examined scatter from the head on the central axis for eight different treatment units. They conclude that, except for one unique collimator design, head scatter originates primarily in the flattening filter and is relatively independent of energy and machine, and is usually less than 5 percent. Significant scatter contributions also originate from wedges and compensators as shown by Huang et al.76 Nordell and Brahmel52 developed an integral expression for the calculation of the absolute yield and angular distribution of photons from any material and for electron energies between 6 and 50 MeV. Good agreement is found with published experimental and theoretical data. The dominant contribution to the absorbed dose outside the useful photon beam is due to phantom or patient scattered photons as determined by a calculation of Nilsson and Brahme.151 Podgorsak et al.170 examined the effects of different atomic number targets and flattening filters on small 10-cm diameter fields 100 cm from the target for energies of 25 MeV. They found that x-ray output on the central axis does not depend significantly on the Z of the target. The dose rate at about 14" determines, after flattening, the dose rate on the axis. An aluminum target gives a more penetrating beam, although high Z targets emit more radiation at large angles. They also found that an aluminum filter hardens the beam and a high Z filter softens it at this energy of 25 MeV. The latter is due to loss of high photon energies by pair production in high Z filters. They recommend a thick aluminum target and flattening filter above 15 MeV for the most penetrating beam.171 Below 15 MeV, a high Z target and a low Z filter are recommended. At 25 MeV, an aluminum flattening filter is 25 cm in length, an impractical size to incorporate in most linac treatment heads. The relatively low melting point of aluminum precludes its use as a target material. Ideally, one wishes flattened fields for all field sizes, at all depths, an impossible requirement because of energy changes and scattering in the phantom. One early 4-MV filter design provided satisfactory flatness at 10-cm depth, but resulted in excessive dose at shallow depths near the edges of
148
CHAPTER 8. TREATMENT BEAM PRODUCTION
large fields.62.86 Invariably, flattening filter choice involves a compromise in order to achieve uniform small and large fields over a range of depths and yet fit into the radiation head. Two flattening filters for a given x-ray energy are provided in some treatment units to optimize beam flattening with a changeover at about the 10 X 10-cm field, or an additional filter may be attached to the accessory ring for assuring large field flatness.22.107 Dual x-ray energy units are equipped with two flattening filters, one for each energy. Removable filters, which affect dosemeter calibration, must be appropriately interlocked. McCall et al.130 examined linac depth dose distributions using a semiempirical analytic depth dose model correlated with experimental measurements. They find that an aluminum filter at 25 MeV produces a great deal more beam hardening on the axis than do nickel and tungsten and consequently, there is a larger variation of photon energy with production angle for aluminum, a significant effect for large field sizes. Aluminum would appear to be a desirable flattener, but the penalty is a significant energy spread across large flattened fields and hence, larger variation in field flatness at depths above and below the design depth. For the Clinac 35, operating at 25-MeV x-ray energy, they recommend a copper target with an iron filter containing a tungsten conical insert for an optimum combination so as to minimize energy variation with angle and restrict beam hardening on the central axis. Other desirable properties of a good target-flattener system are that the flattener should not become too radioactive in operation, and that neutron production be minimized.130 Thus, copper filters are not commonly used above 10 MeV since the gamma ray dose rate from 9.76 min T% 62Cu activation becomes very high. Iron, on the other hand, has essentially the same absorption properties as copper, and the induced activity is much weaker. Flattening filters may absorb 50-90 percent of the central axis photon intensity. They, as well as wedge filters, increase scatter, most noticeably outside the geometric confines of the beam.76 Brahme et a1.,27 recommend scanned photon beams at very high energies (25-50 MeV) pointing out that the flatness and intensity problems are greatly alleviated. Hence, some radiation shielding problems (including those for neutrons) may be as much as two to five times less for 20 to 50-MeV scanned x-ray beams than for a heavily filtered unscanned beam at such energies. The spectral change problem associated with flattening filters was cited earlier by Hansen, et al.65 who identified its importance for a 4-MV linac beam. Later, Larsen et al.113 developed a calculative program for filter design and applied it to a 4-MV beam. Their program summarizes the primary and scattered components in an iterative manner to fit the dose profile. Flock and Shragge54 developed a semianalytical method for the design of an improved Therac 20 flattening filter. Jones95 points out that the dominant effect is selective hardening of the beam by the flattening filter and, that for thin targets, the effective energy decreases with distance off axis.
Although an x-ray beam can be within flatness specification at 10-cm depth, it may be significantly unflat at the depth of dose maximum. Constantinou and Sternick44 found a 19 percent increase in the off-axis "horns" of a 6-MV unit, which was reduced to 8.5 percent by increasing the beam energy. A 1 percent increase in beam energy resulted in approximately a 1 percent reduction in the horns at the depth of dose maximum and is associated with the angular distribution of bremsstrahlung as a function of beam energy. Hanson et al.66,67 measured the off-axis quality change for 4- to 10-MV beams and suggested a technique to correct for the effect in treatment planning calculations. Lutz and Larsenl21 looked at the effect of flattening filter design on quality variations within an 8-MV primary x-ray beam. They found that the single material flattening filter supplied by the manufacturer introduces most of the quality variation. By adding a hardening filter or employing a composite brass and lead filter, the quality variation was reduced by one-half, but output is reduced by 25 percent. Flattening an x-ray field involves a compromise between penetrability, uniformity, and output. Contemporary machines employ energy interlocks whose limits place constraints on field unflatness. Naylor and Chiverallsl48 examined the variations in x-ray beam flatness and calibration as a function of gantry angle over a period of time for an 8-MV unit equipped with a 90' bending magnet. They conclude 'that such variations are confined to a few percent when averaged over 4-week periods, an interim pertinent to treatment. In one treatment unit, Dale found that an anomalous asymmetry with gantry angle was caused by the leads of the electron chamber sagging into the edge of the useful beam.45 Padikal et a1.,156 describe a method for assessing the stability of symmetry for gantry angle rotation. The effect of the earth's magnetic field on the electron beam in the linac can change with gantry rotation. The heavy iron frame cast in the floor and used for mounting the linac can produce significant distortions in the earth's field and changes in fringing fields of bend and solenoid magnets as a function of gantry angle of rotation. Sutherlandlgl notes that dose monitoring methods employing circular chambers bias their output for unflat fields so as to affect calibration factors by as much as 2 3 percent. In an early study, Naylor and Williamsl49 call attention to the need for frequent symmetry,dose monitor, and beam energy checks of linac treatment units. Martell et al.,129 describe a flatness monitor that can be used to quickly and accurately check beam flatness and calibrations. It consists of a 7 X 7 matrix of ion chambers, a sampling multiplexer, and computer hardware and software to provide a hard copy printout. An earlier matrix dosemeter for the uniformity of high energy x-ray beams is described by Johns et al.89
X-RAY SCANNING SYSTEM At x-ray energies above 25 MeV the x-ray lobe becomes so narrow that use of a conventional flattening filter for large fields becomes undesirable for a number of reasons. The pri-
149
X-RAY THERAPY
mary radiation photon absorption process in the x-ray target at such energies results in positron~lectronpair production, with consequent reduction in mean photon energy in the central portions of the field. There is not enough space available in the radiation head of 360" isocentric linacs at these high energies to permit use of low Z materials in the central portion of the flattening filter to avoid this radius-dependent degradation of the x-ray spectrum (as studied by Podgorsak et al.).l70 Also, the mean displacement of the beam at the x-ray target would have to be controlled still hore precisely at these high energies if a flattening filter were used, because of its greater centrallperiphera1 attenuation ratio and because of greater geometric magnification if the filter length were increased by placing its tip closer to the x-ray target than in conventional designs. Figure 8-8 shows a magnet system design proposed by Brahme et al.27 for producing a scanned x-ray beam. The narrow energy spread (- 0.1 percent) horizontal beam from a 50-MeV racetrack microtron is bent by a 90" bend magnet onto an x-ray target in a treatment gantry. A radial scanning magnet is located ahead of the 90" bend magnet. An electron ray diverging radially from the center of this radial scanning magnet reconverge at the center of a transverse scanning magnet located after the 90" bend magnet and just before the x-ray target. Although the angular scanning enlarges the beam spot on the x-ray target, the x-ray lobe is so narrow at 50 MeV that at the normal treatment distance it appears to originate from the electron beam radialltransverse convergence/divergence point in the middle of the transverse scanning magnet ahead of the x-ray target. A raster scan or a spiral scan in the angle of the x-ray beam lobe can be generated by appropriate control of timing and amplitude of the currents in the two scan magnets. Scanning of the x-ray lobe angle can produce a relatively flat dose distribution with only a thin x-ray filter needed to
correct small nonuniforrnities. Thus, the effective energy of the x-ray spectrum remains high and uniform over the treatment field, even for large fields. At 50 MeV, the full-width of the unflattened x-ray lobe at 50 percent intensity is about 8 cm at 100-cmnormal treatment distance. Overscanning onto the jaws of the beam limiting system is required to obtain adequate overlap of the individual beam lobe pulses for uniform dose distribution out to the edges of the maximum field size. Thus, some of the x-ray power that would be lost from the central portion of the x-ray lobe with a conventional flattening filter is lost instead outside the periphery of the maximum field. X-ray beam scanning could be usefully applied at energies lower than 50 MeV as well. For example, at 20 MeV the diameter of the unflattened lobe at 50 percent intensity is about 20 cm so that a 40 X 40-cm maximum field size would require scanning the x-ray lobe axis over approximately a 60 X 60cm area. The maximum treatment field area would be approximately 44 percent of the x-ray scan area. If a flattening filter were used instead, the flattened field intensity would be approximately 16 percent of the unflattened lobe maximum at 50 MeV. Thus, for a given dose rate the required beam current and power on the x-ray target and consequent stray x-ray and neutron emission would be reduced by a significant factor (a factor of 2V2) even at 20 MeV by use of scanning. A single 90" nonachromatic bend magnet can be used in the system of Figure 8-8 because of the narrow energy spread of the microtron. For linacs, either a major reduction of the energy spread or an achromatic bend magnet would be required. Radial scanning ahead of the bend magnet over the approximately 2300 milliradians needed to adequately overscan the treatment field would far exceed the acceptance of achromatic bend magnets. Thus, for linacs, it is more practical to scan in both radial and transverse planes at a point after the
-
Bending Magnet Electron Beam
n
Scanning Magnets
11
FIGURE 8-8 . Cross sections of magnet system for scanning the angle of arrival of the beam at the x-ray target (from Ref. 28). (a) The radial plane, (b) the transverse plane, and (c) the adopted scanning magnet system.
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CHAPTER 8. TREATMENT BEAM PRODUCTION
bend. Bensussan et al.17 proposed such a technique for x-ray scanning of a linac beam, employing a four-pole magnet.
SCANNED BEAM DOSIMETRY Patient dosimetry and beam monitoring is more difficult with scanned electron and scanned x-ray beams. Because of the narrow radiation lobe and its short pulse length, the instantaneous ionization intensity is high in the core of the radiation lobe in the ionization chamber of the machine dose monitoring system. Collection efficiency can be significantly impaired, requiring correction for resulting nonlinearity.141 Measurements of the dose distribution in a phantom can be made with integrating techniques, such as by using film, but use of scanning probe type detectors is more complex and may involve the linac pulse repetition frequency. More complex interlocking is required to ensure that dose uniformity over the field is not affected by missed accelerator pulses (see also Chap. 9).
CONTAMINATION OF RADIATION BEAM A treatment beam of x-rays or electrons contains contaminants, both inside and outside the useful beam intended to encompass the tumor. These contaminants are electrons and photons as well as neutrons, which are considered separately in the following section. The contaminants modify the intended dose distribution pattern deposited inside the useful beam as well as exposing tissues outside to unnecessary and, perhaps, harmful radiations. Fraass et a155 pointed out the wide variation in thresholds of deleterious dose-effect phenomena ranging from less than 5 to 200 cGy or about 0.1 to 3 percent of the prescribed doses. Contaminant limits have been proposed by the IEC85 in terms of leakage radiation through beam limiting devices, surface dose during x-ray irradiation and stray radiation during electron irradiation. The major contaminants of megavoltage x-ray beams (electrons, scattered photons, and leakage radiation), have received the most study. Here the important effect within the useful beam is the reduction in skin sparing which results from increasing skin and build-up dose and a shift of the dose maximum towards the surface as the field size is increased. Efforts to ameliorate this effect have been successful and many investigators have contributed to the study and reduction of these surface dose phenomena~19,23,74,76,96,103,111,118,124,125,161,166,194,2M
The loss of skin sparing is due primarily to electrons as shown by Biggs and Ling19 using a sweeping magnet placed just below the radiation head of a megavoltage linac. This loss can be partially regained with transparent lead-loaded acrylic filters.118 Huang et al.74 report a 3 to 13 percent reduction in surface dose compared with pure acrylic, which is often employed for supporting field blocks. Build-up data for compari-
son purposes must be obtained with operational plastic shadow trays in place. The x-ray flattening filter has been suggested by experiments of Mackie and Scrimgerl24 as the main source of contaminant electrons. Monte Carlo studies confirm these findings.1517166The main sources of scattered photons are the primary collimator for 6 MV and the flattening filter for 2 1-MV bremsstahlung sources.151 The contaminants outside the useful beam of x-rays involve lower doses but larger volumes and may expose radiation sensitive entities such as the fetus or the lens of the eye. These peripheral doses have been the subject of a number of studies.55,63,@,96,186,189Electrons may contribute as much as 20 percent of the dose maximum up to the depth of dose maximum. However, other components, phantom or patient, and collimator scatter and leakage transmitted through the radiation head, are dominant at greater depths. Fraass et a1.55 find inphantom scatter and leakage dose to be roughly equal and varying with the field size, while Greenea and Kase et al.96 identify collimator scatter as a major source. Leakage radiation becomes dominant at large distances from the electron beam axis although dose from leakage is small and the cost in money and weight to reduce this dose component further than the 0.1 percent limit guideline may not be warranted. Wedges and blocks contribute dose outside a treatment field. Sherazi and Kase186 found that wedges can increase such doses by a factor of 2 to 4 from the unwedged value but the effect of adding blocks is much smaller and generally less than a factor of 2. Electron treatment beams are contaminated by bremsstrahlung, which dominate the depth dose beyond the practical electron ranges. This dose increases with energy but can be minimized by attention to radiation head design. The surface dose of electron beams increases with field size. A Monte Carlo study by Udalel98 relates this effect to the adjustable x-ray collimators that shield the phantom or patient from secondary electrons for smaller fields, and to the electron applicators that increasingly filter the beam as they get narrower for smaller fields, thereby raising the mean energy of the electrons. Photon contamination in electron arc therapy has been studied by Pla et al.169 It is a function of arc angle, field width and isocenter depth and under certain conditions can amount to a large fraction (up to 50 percent) of the prescribed electron dose.
NEUTRON LEAKAGE AND RADIOACTIVATION A significant number of neutrons are produced by high energy x-ray beams. The neutrons and the radioactivation that they may induce have been the subject of a number of studieS~2,7,18,51,69,70,71,109,110,I30-133,1,147,153,157,183,200 For most relevant materials, the neutron production threshold occurs at 8 to
REFERENCES
10 MeV, rises rapidly, and then plateaus above 20-MeV photon energy. Neutrons that originate in the primary collimator, target, and flattening filter contaminate the useful beam. Others are filtered through the radiation head, a few are generated in the patient (see Chap. 14,p 252 and many are multiply scattered by barriers comprising the treatment room, which themselves are a source of neutrons. These neutrons, together with x-ray leakage, can expose the patient as well as those outside the treatment room. McCall and Swansonl3l provide a thorough description of neutron sources and their characteristics originating in linac radiation heads. They find that high Z shielding materials do not significantly alter the neutron fluence, but do substantially reduce the average energy of the transmitted spectrum and, hence, their contribution to dose. The principal source of neutrons at 25 MeV is the primary collimator, which contributes one-half or more of the fluence, often followed by the target and flattening filter in that order. Depending on the material, neutrons can induce radioactivity in the radiation head and patient support components of the linac, as well as the treatment room barriers. Such radioactivity could constitute an appreciable source of exposure, particularly for the technologists. With respect to neutron production, a recent measurement showed that a tungsten flattening filter, five radiation lengths thick, produced two and one-half times as many neutrons as an equivalent steel filter at 25 MV,lIO although ratios as high as seven times have been predicted.131 Muller-Runkel et a1.140 discuss a maze designed to reduce the flux of neutrons at the door to permissible levels for a Therac 20 accelerator. Neutron measurements in medical accelerator rooms have been described by La Riviere for a 24-MV linac.1" Uwamino et a1.200 describe extensive neutron measurements and calculations for a facility employing a Scanditronix MM22 microtron. A comprehensive report describes neutron measurement methods around high energy x-ray radiotherapy machines.' Examination of neutron biological effectiveness has led to a scrutiny of neutron protection requirements by regulatory agencies. This scrutiny and the associated need to verify neutron production data, as well as to establish appropriate measurement, calculation and dose reduction techniques were the subject of a conference.69 A consensus appears to have developed wherein proposed neutron leakage requirements can be more easily met in contemporary high energy linacs.70J3~.~47 Often, the shielding provided for x-ray leakage is sufficient for neutron leakage as well, and satisfies the total leakage limitation of 0.1 percent of primary beam measured in centigray. Rawlinson and Johns178 note that the energy imparted outside the useful beam due to x-ray scatter occurring in the patient can be more than 20 times greater than the energy due to radiation leakage. Some simplification in assessment have resulted from defining two pertinent measurement surfaces for protection purposes.85 One is a plane circular surface, that is, "patient plane" of radius 2 m centered on and perpendicular to the axis of the beam at the normal treatment distance. This relates to patient protection. A second cylinderlike complex surface is defined by all points at 1 m from the path of the electrons
151
between the electron gun and the target or electron window. This relates to room shielding. The report of an AAPM Task Group on neutrons from high energy x-ray medical accelerators provides a carefully reasoned, quantitative analysis of the neutron problem together with recommendations regarding risk to the radiotherapy patient.147The report concludes that the risk of carcinogenesis is extremely small, and the implementation of more restrictive regulation is unnecessary and would be counterproductive.
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Dose Monitoring and Beam Stabilization
Radiation therapy involves the identification, adoption, and implementation of an optimal treatment plan for the individual patient. Such a plan includes a treatment prescription specifying the individual field sizes, their orientation, and their individual contribution to the composite dose distribution in the target volume. It includes the fractionation scheme (sequence of daily treatments) for delivering the prescribed total dose, as well as provision for planned changes in fields during the course of therapy. In addition, the treatment plan may include the use of beam modifying accessories such as wedge filters, compensators, shadow blocks, or combinations of these. The safe, quantitative, and accurate delivery of the individual field dose portions of the treatment prescription is the central function of the dose monitoring and beam stabilization system. Radiation treatment beams must be precisely directed and accessories accurately positioned in order to achieve maximum clinical benefit. See Fig. 13-4 for an example treatment prescription. The dose monitoring system incorporates a transmission ionization chamber. This chamber is located in the radiation head of the linac and samples the treatment beam (see Fig. 8-1). Electrical signals from it are used to monitor and control the treatment beam. Various aspects of electron dosimetry including monitoring have been described in two International Commission on Radiation Units and Measurements (ICRU) reports.16.20 Ageneral reference on the physics of electron beam therapy is that of Klevenhagen.28 The dosimetry of x-ray beams is described in ICRU Report Nos. 23 and 24,17918 and NCRP Report No. 69.41 The dosimetry of pulsed radiation, especially relevant to electron linacs, is treated in ICRU Report No. 34.19 IEC document 601-2-1, specifies radiation safety performance standards and test requirements for dose monitoring systems.21 Patient dosimetry and beam monitoring is more difficult with scanned electron and scanned x-ray beams. Because of the narrow radiation lobe and its short pulse length, the instantaneous ionization intensity is high in the core of the radiation lobe in the ionization chamber of the machine dose monitoring system. Collection efficiency can be significantly impaired, requiring correction for resulting nonlinearity. Measurements of the dose distribution in a phantom can be made with inte-
grating techniques, such as by using film, but use of scanning probe type detectors is more complex and one must account for the linac pulse repetition frequency. More complex interlocking is required to ensure that dose uniformity over the field is not affected by missed accelerator pulses.
IONIZATION CHAMBER The transmission ionization chamber of a contemporary high energy linac is constructed of several plates or electrodes, whose areas may be divided into sectors so as to serve two different monitoring purposes. 1. Dose rate and integral dose of the x-ray and electron treatment beams. 2. Angular and radial (positional) distribution of the radiation in the treatment In many accelerators, the resultant electrical signals are also used in automatic feedback circuits to steer the electron beam through the accelerator and bending magnet onto the target (or scatterer) in order to ensure treatment beam flatness and symmetry. How these needs are satisfied in one representative treatment unit, the Varian Clinac 18, is described below and illustrated in Figure 9-1, which is a simplified diagram of the ion chamber, dosimetry, and beam steering system. Additional details of this35 and othefi.48.52 monitor chamber designs are described in the patent literature. Ionization chamber position misalignments can significantly alter their response. Hence, precise repositioning is critical where two independent moveable chambers are employed for x-ray and electron beam monitoring, especially where heavy flattening filters are also moved. Figure 9-2 illustrates typical electron paths, between gun cathode and x-ray target and the beam steering system for the Clinac 18. Figure 9-3 is the logic diagram of the dual channel dosimetry system of the Clinac 18 as described herein. The transmission ionization chamber, shown in the radiation head diagram of Figure 8-1, subtends the entire useful
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CHAPTER 9. DOSE MONITORING AND BEAM STABILIZATION
Position Steering Coils
Angle Steering Coils
FIGURE 9-1 . Five-electrode ionization chamber with simplified block diagram of dosimetry and beam steering system. The radial and transverse coordinate planes of the bending magnet orbit are identified in the upper left. The radial angle steering coils are actually located in the bending magnet, as shown in more detail in Figure 9-2.
beam and provides two independent outputs for the dual dosimetry monitor described in the section that follows. Located just below the x-ray flattening filter or electron scattering foil, it monitors the emergent flattened x-ray beam or scattered electron beam. One design of this ionization chamber, shown in more detail in Figure 9-1, consists of three polarizing plates and two intervening collecting plates- Each of the latter is
divided into four collecting sectors, with each sector defining a distinct laminar collecting volume. The parallel-plate construction of the ion chamber allows monitoring the entire useful beam emerging from the primary collimator and permits close spacing of the plates, ensuring good ion collection at moderate voltages. Treatment beam x-rays and electrons eject secondary electrons from the plates into the gas between the plates,
159
IONIZATION CHAMBER
3rd Polepiece
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/
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I :
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FIGURE 9-2
Gun Anode
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' 1st Polepiece
Transverse Angle Steering Coil
Input Collimator
. Accelerator, bean1 steering, bend magnet, and typical electron paths for Clinac 18 treatment unit. (Courtesy of Varian)
I
I
COLLECTOR 1
I
INTEGRATOR
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DOSE RATE
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,
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I
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FIGURE 9-3 . Dual dosinletry logic for ionization chamber systenl illustrated in Figure 9-1. (Courtesy of Varian.)
ionizing it, with some ionizing of the gas also caused directly by the treatment beam. An energetic electron will typically produce about 60 ion pairs per centimeter of path in air at normal temperature and pressure (NTP). The positive and negative ions drift in opposite directions towards the two plates under the influence of the electric field between the plates. The electrons produced in the gas are attracted to the electrode and + ions to the - electrode, the fraction arriving being dependent on the collecting voltage, plate spacing, dose per pulse, and the particular gas employed and its density. Some ions recombine, either from the initial pair created (initial recombination) or from close encounters with ions of opposite polarity during the collection process (volume recombination). The frequency of the latter depends on the density of ions created and the collecting field strength. Saturation is defined as collection of all of the ions produced. To avoid significant nonlinearities at operational dose rates, an ionization chamber is usually operated very near saturation. A small gap and high collecting voltage minimize this recombination loss, which is nonlinear and a function of dose rate. A 500-V polarizing voltage is used with a plate spacing of 1 mm. The two inner D-like sectors provide signals for both dosimetry and angle steering and the two outer arclike sectors for displacement steering. The dosimetry system monitors and displays readings related to the quantity (dose), dose rate and uniformity (symmetry) of the useful beam of radiation. In one construction technique, the collecting plate sectors are formed by vacuum deposition of a thin metallic coating on defined areas of an insulating lamina of mica. Additional grounded metallic coatings, not shown in Figure 9-1, surround
+
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CHAPTER 9. DOSE MONITORING AND BEAM STABILIZATION
the collecting areas and serve as guard rings to minimize leakage currents over the insulation. Ionization chambers may be sealed to the outside air, making them free of the need for temperature and pressure corrections provided that the pressure diaphragms are not used as electrodes. Hrejsa et al.,ls describe a temperature-pressure compensation system for an unsealed ionization chamber.By its use, observed fluctuations of the dose calibration per monitor unit were reduced from 3 to 6 percent to approximately 20.5 percent over extended periods of time. The thin exit and entrance windows of sealed, parallel plate transmission ion chambers may flex with ambient pressure changes altering slightly the mass of the central volume of gas being monitored and, hence, the calibration. Sealed ion chambers are preferably filled to a positive pressure above atmospheric pressure. If chambers leak, or are intentionally vented to the atmosphere, their sensitivity will also shift with ambient pressure and temperature changes.5.47 Assessment of such variations in monitoring ion chamber sensitivity is difficult under routine treatment unit operation. An improved testing procedure for sealed, parallel plate monitor chambers has been developed by Kehoe and Barnard." By placing the chamber in a testing jig, they are able to simulate and exaggerate ambient pressure variations and detect leaks quickly. Sutherland49 recommends collecting ion current for dosimetry only from the central part of the ion chamber, guarding against the effect of the more unstable peripheral rays, which may unduly bias ion chamber response. A relatively new material for ion chamber plates is Kapton," a synthetic polyimide, thin film. It is extremely resistant to radiation damage, compared to the other insulating materials used in ion chamber construction. It can be readily plated with copper to which leads can be soft soldered. A chamber of this construction has significantly fewer grams per square centimeter in the electron beam than a mica plate chamber. While not an important consideration for x-ray therapy, it allows the discrete electron scatterers to more completely control scattering for electron therapy. Kapton film is very strong, not subject to adverse effects of heat, and may be stretched taut. However, some sealed, nitrogen filled, Kapton chambers developed troublesome calibration changes after several months of routine use. Dry-air-filled Kapton chambers appear to be free of this anomaly.
INTEGRATED DOSE AND DOSE RATE Contemporary linacs employ two independent integrated dose monitors, a dose rate monitor and a backup timer, which should be preset to preclude excessive dose delivery by restricting treatment "beam-on" time. As shown in Figure 9-1, the dose monitors are fed by the inner D-like sectors of the two collector plates. The dose rate monitor derives its signal from both the *Kapton, Du Pont de Nemours & Company, Wilmington, DE 19898.
inner D-like and outer arclike sectors of the upper collecting plate. Sutherland49 early recommended that integrated dose be based on monitoring only the central portion of the field. Most treatment fields are small, and he found that monitoring the large area of the entire beam profile can result in axial calibration errors as large as 3 percent. An ion chamber incorporating a circular area, centered on the beam axis, has been described by Boux6 and by Stieber.48 For safety, two independent dose monitors are needed. The electron beam center line (central orbit) through the linac structure and 270" bending magnet is shown in Figure 9-1. This center line determines the reference orientation of the semicircular plates with respect to the radial and transverse coordinate planes of the beam as established by the bending magnet. As shown, the upper collecting plate of Figure 9-1 is concerned with signals pertinent to the radial plane of the bend magnet, the lower collecting plate to the transverse plane signals. Semicircular electrodes, A and B, are oriented to provide signals related to the radial plane. Their current signals are first converted to voltage and amplified via A and A2, then summed via A3 to provide a console indication of dose rate and of integrated dose via integrator no.1 for the DOSE 1 channel as shown. Similarly, semicircular electrodes C and D are oriented to provide signals related to the transverse plane. They feed the DOSE 2 channel via amplifiers A5 and A6, summing amplifier A, and integrator no. 2. The two integrateddose channels are completely independent; either can terminate the preset exposure or if the second channel lags the first by more than 25 monitor units. Hence, if the treatment is not terminated normally by the first dose channel on completion of the preset monitor units, the second dose channel will terminate the treatment 25 monitor units (MU) later or at + 10percent, whichever is lower (as required by some State Laws). The dual 500-V power supply furnishes independent collecting voltages for the two ion chambers. Dose monitors are calibrated on site so that a monitor unit (MU) read out on the control console is uniquely and unambiguously related to the absorbed dose in a phantom under prescribed conditions. For example, 1 MU may equal 1 cGy for a 10 X 10-cm field on axis at 100-cm SSD and 10 cm depth (or at D,,,) in a water phantom. Ionization chambers may exhibit nonlinearities as a function of dose rate if there is significant ion recombination.54 In this case the ions that are formed in the gas, usually air, between the collecting plates recombine before reaching the collecting electrodes. Ion collection is more difficult in the radiation beam of linacs where the pulsed nature of the beam results in high ion density prior to collection. Hence, large collecting voltages are employed to provide high collecting E-fields for overcoming recombination. In well-designed systems and under normal operation, ion collection efficiencies over 99 percent are achievable with x-rays, and somewhat less with electrons, particularly at lower (4-6 MeV) electron energies. The dosimetry of pulsed radiation beams and a simple two-voltage technique for assessing collection efficiency in pulsed beams has been described.19 Several radiotherapy linacs now incorporate
+
'
FIELD UNIFORMITY CONTROL
magnetically swept rather than scattered electron beams, which present different design and additional ion collection efficiency problems.8.36.ss Conere and Boag9 extended the two-voltage assessment technique cited above to such beams. Johnsen and La Riviere23 examined the collection efficiency of commercial thimble ionization chambers in pulsed x-ray and electron beams. Using a 0.6-cm3 Farmer chamber they measure collection efficiencies of above 99 percent and as low as 90 percent for exposures of 0.04 Rlpulse for routine radiotherapy and 0.5 Rlpulse for scanned electron beams respectively. O'Brien et al.43 calculated the collection efficiency for this chamber as between 14 percent and 10 percent when exposed to the abnormal dose per pulse of between 1 and 2 Gy. Pillar et al.46 describe an independent dose monitor totalizer unit that is especially useful for patients being treated at extended distances when it is necessary to reset the control console dosemeter several times. Dependence of integrated dose on dose rate can stem from nonlinearities of ion chamber electronics as well as from ion recombination. Cheng and Kubo7 found that a defective circuit board resulted in a 20 percent decrease in integrated dose for a dose rate increase from 100 to 500 MUlmin. Cox et al.10 found amplifier saturation a problem in developing a total skin electron therapy (TSET) technique, which required a very high beam intensity at the location of the dual ion chamber monitors in the radiation head. By improving amplifier linearity, increasing the collector voltage, and requiring the primary dosemeter to be connected to the most distant of the dual chambers from the scatterer, satisfactory monitoring was assured using existing monitor chambers in the radiation head. The initialized use of shall below indicates a mandatory requirement. Dual dosimetry monitors, which provide independent exposure limit backup, are recommended by safety organizations.21 Malfunctioning of one dose monitoring system shall not affect the correct functioning of the other system. Failure of an element which could affect the correct function of either dose monitoring system shall terminate irradiation. In the case of separate power supplies, failure of the power supplies of either system shall terminate irradiation. If the performance of the detectors is dependent on hermetic sealing, then they shall be independently sealed. The two dose monitoring systems shall be arranged either as a redundant dose monitoring combination or as a primarylsecondary dose monitoring combination. In the case of a redundant dose monitoring combination, both systems shall be capable of a given performance. In the case of a primarylsecondary dose monitoring combination, at least the primary dose monitoring system shall be capable of this performance. This performance shall be maintained up to absorbed dose rates of twice the specified maximum. The secondary dose monitoring system shall be set to terminate irradiation when the preselected number of dose monitor units has been exceeded by not more
161
than 15 percent if a percentage margin is used or not more than the equivalent of 40 cGy of absorbed dose if a fixed margin is used. The digital logic of many dosemeter and control systems facilitates the development of rapid automatic self-test regimens of dosemeter and other circuits prior to each treatment exposure. In such self-test systems, a known current substitutes for the ion current during a short fixed time, producing a pseudodose integrand whose value assesses the correct function of the dose integrating circuit, but not the radiation response of the ionization chambers. Under computer control, the dose integrating circuit may be tested in this manner prior to each treatment exposure. Figure 9-3 illustrates the logic diagram for a dual channel dosimetry system wherein the backup channel will terminate treatment at a dose 40 MU higher than the dose set on the thumbwheel. The self-test calibrate current is introduced in each ionization chamber collector plate lead. The interlock also terminates the treatment if either of the preset radial or transverse symmetry limits are exceeded. Digital circuits can be very reliable, are less subject to failure, are easier to design, and are increasingly incorporated in monitor and control systems. They exhibit freedom from drift and need fewer adjustable controls than analog circuits. They can serve to implement software solutions to problems, often as test regimes. It seems likely that microprocessor and firmware will be increasingly incorporated into radiotherapy linacs as in other computer applications. This allows the flexibility of program development in software and, once agreed, the "freezing" of the program in inexpensive read only memory (ROM). However, the cost of software development is often many times the cost of the hardware development. Additional aspects of digital logic, monitor and control systems as well as computer usage are presented in Chap. 10.
FIELD UNIFORMITY CONTROL The field uniformity control system is designed to ensure that the treatment field is symmetrical in both the radial and transverse directions and uniform to within the stated specification of the manufacturer. This entails control of the beam energy and the accurate placement of the beam on the axis of the target-flattening filter for x-ray therapy or the scattering foil system for electron therapy.6 This placement can be accomplished by the use of steering coils located along and around the accelerating structure, as well as at the bend magnet. By energizing them from appropriate ionization chamber signals, beam control in both radial and transverse directions is provided. Figures 9-1 and 9-2 illustrate the beam steering system, which provides field uniformity in the Clinac 18. The radial and transverse planes intersect along the beam center line and
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CHAPTER 9. DOSE MONITORING AND BEAM STABILIZATION
the axis of the x-ray flattening filter (see Fig. 8-1). As shown in Figure 9-1, two groups of four steering coils are used in a servofeedback system, using signals from the ion chamber sectors to control and limit the divergence (angle) and displacement (position) of the electron beam from the axis in the radial and transverse directions. Both groups of steering coils provide small angular deflections of the electron beam. Adetailed view of the Clinac 18 accelerator and beam steering and bending magnet system is illustrated in Figure 9-2. The transverse angle steering coils shown on the left in Figure 9-1 are located at the entrant image aperture of the 270" bending magnet, as illustrated in Figure 9-2. The exact location of such coils may vary, depending on the design of the overall beam transport system, but magnet optics dictate their location in a particular design. In an achromatic magnet, an entry trajectory angle at the entrant image plane results in the same exit angle at the object image plane without producing a change in displacement at the image plane. The radial angle steering coils are located at the second polepieces of the three-dipole bending magnet as shown in Figure 9-2. The position steering coils, shown on the right in Figure 9-1, are located at a distance from the bending magnet entrant aperture and at the output end of the accelerating structure as illustrated in more detail in Figure 9-2. As shown in Figures 9-1 and 9-2, two of the beam position steering coils are connected to control the radial position, and two, the transverse position of the electron beam from the beam center line axis. The principal effect of their small angular deflection is to provide a lateral displacement (position) correction at the bending magnet entrant aperture, which is at or near the beam collimator. The small angular error introduced is then corrected by the angle servo. These steering coil groups are energized by error signals generated if the electron beam strikes the target, or scattering foil, at an angle or at a position that produces an asymmetrical x-ray or electron beam, as detected by the two monitor chambers and as shown in Figure 7-16b and c, respectively. In the Clinac 18, a third group of four coils, shown in Figure 9-2, is also positioned around the beam at the buncher (input) end of the linac structure. These unsewoed buncher coils also control beam position in the radial and transverse planes, steering electron beams of all energies leaving the gun onto the center line of the microwave accelerator structure in the first few cavities. Signals from peripheral plates E and F are amplified by amplifiers A9 and Ale. Their difference signal (E-F) from A,, feeds a "sample-and-hold" circuit, which provides a dc signal to the radial position steering coils. The radial position steering coils control the radial component of lateral displacement of the electron beam with respect to the flattening filter axis, as shown in Figures 9-1 and 7-16 c. Amplified signals from semicircular plates A and B are subtracted in difference amplifier A,, which feeds a sample-and-hold circuit which, in turn, provides a dc signal to the radial angle steering coils. The radial angle steering coils control the radial component of angular divergence of the electron beam with respect to the x-ray target and field flattening filter axis (see Figs. 9-1 and 7-16 b). The angle and
position asymmetry signals from amplifiers A4 and A l , and their associated sample-and-hold circuits, are compared to provide a visual display of radial plane beam asymmetry at the console. They are set to provide an operational radial asymmetry limit beyond which the beam is turned off. In a similar manner, as shown in Figure 9-1, but not described here, the transverse position and transverse steering coils are connected to amplifiers and ion chamber sectors so as to semocontrol, display, and limit the beam asymmetry in the transverse plane. The symmetry meter (see Fig. 9-I), can be switched to indicate either radial or transverse symmetry. The symmetry ineter and the associated steering coils are connected to use all signals from the collecting electrodes. Asymmetry interlock interrupts the beam if asymmetry exceeds a preset value (e.g., 2 percent). The beam position and beam angle steering amplifiers are each provided with six programmable groups of gain and balance controls corresponding to the one x-ray and five electron modes of operation of the Clinac 18.The associated steering interlocks are verified during Quality Assurance (QA) procedures. The beam position signals are not used in the electron mode because the electron beam is not widely spread at the ion chamber and the precise steering associated with high energy, x-ray flattening filters is unnecessary. Here, the signals from the small outer plates E, F and G, Hare small and provide little useful information. Quality assurance procedures are particularly important if a broad beam illuminates the primary collimator and steering interlocks are not employed.
MONITORING AND CONTROL OF MULTIMODALITY TREATMENT UNITS Multimodality treatment units involve significant beam monitoring and stabilization requirements, and present a hazard that is not found in single modality units. The radiation safety hazard of high abnormal electron beam currents in an early dual modality treatment unit was identified many years ago.25 A particular combination of electronic and mechanical malfunctions, combined with failure of the software to respond properly to an operator action, appears to have been responsible for several overexposures in 1985-1987, where a large electron bearn current emerged as an unscattered, unscanned (and almost unmonitored) beam from the radiation head. In normal x-ray therapy, the electron beam is intercepted by the x-ray target or by the x-ray flattening filter if the target is retracted erroneously. In normal electron therapy, the electron beam is spread out by a scattering foil(s) or scanned over the treatment field magnetically. In normal electron therapy mode operation, the beam current through the electron window is of the order of 111000 of the beam current at the x-ray target for xray therapy mode. For example, at 4 Gylmin at 100 cm in the 6MV x-ray mode, the average beam current at the x-ray target in one linac treatment unit is of the order of 100 PA. At the same
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MONITORING AND CONTROL OF MULTIMODALITY TREATMENT UNITS
dose rate in 6-MeV electron mode, the average beam current at the electron window is of the order of 0.1 pA. Similarly, an average beam current of 20 pA is typical for 25-MV x-ray therapy, and approximately 0.02 p A for 25-MeV electron therapy. Typical beam currents at the accelerator vacuum window are given in Table 9-1 for specified x-ray and electron dose rates as a function of energy. Through hardware and/or software failure, a large electron beam current intended for x-ray operation can emerge without being intercepted by the x-ray target andlor x-ray flattening filter. Even with the scattering foil(s) in place or scanning operable, an estimated dose comparable to a typical 2-Gy dose fraction can be delivered to the patient in a time of the order of 0.03 seconds at a dose rate of 4000 Gylmin. This is far too short a time for operator reaction, so safety protection is totally dependent on fast monitoring and radiation terminating electronics. The hazard is exacerbated if it is possible for the electron scattering foil@)or scanning system to fail at the same time as the failure described above. Recently, O'Brien et a1.42.43 measured a dose of l to 2 Gy per pulse from 25-MeV electrons at the normal treatment distance for such abnormal operating conditions. The electron beam lobe will then be more forward peaked at the patient, and the electron dose rate can be of the order of 15,000 Gylmin and deliver a normal 2-Gy dose fraction to a small volume of tissue in only one pulse of the accelerator. Clearly, the safety electronics must include the provision to monitor the radiation beam on a pulse-by-pulse basis and to terminate radiation within one interpulse period, that is a time of the order of 0.002 seconds. The experiments of O'Brien et al.,43 pertinent to this hazard, also demonstrate the severe limitations of many of the commonly used calibration standard dosemeters when they are used outside of the normal dose and dose rate ranges encountered in a radiotherapy department. The quantitative values presented above will vary from one treatment unit to another, depending on the reference dose rate, the pulse repetition frequency, the electron energy and its TABLE 9-1
-
energy spectrum, the design of the x-ray target-filter system, and the electron scattering or scanning system. However, they are believed representative of contemporary designs. Protection of the patient against hazards from unwanted or excessive radiation are reviewed in an IEC document.21 One approach to the overexposure hazard where very high dose rates are possible (more than 10 times normal maximum rates), is to incorporate an additional independent monitor whose sole function is to monitor and interlock the radiation beam on a pulse-by-pulse basis. It has been recommended that this detector be operated on a different physical principle than those already employed.21Possible candidates include the electron secondary emission monitor, as well as electrostatic and electromagnetic induction monitors.l4,24.50 Significant experience with such devices may be found in many high energy particle physics laboratories. Frequently, operator error andlor procedural deficiencies are involved in radiation accidents. Suggestions for improvements in these aspects have been presented.26 Loyd et al.32 assessed the dose delivery error in a dual photon energy, computer-controlled Philips SL 25 linac. This unit has a complex collimation and beam delivery system, is magnetron powered and provides eight electron energies from 4 to 20 MeV as well as 6 and 15-MV x-rays (see Fig. 8-2). It is equipped with a motor-driven nominal 60" wedge for providing variable wedge angle fields, and independent upper collimator jaws for providing asymmetric fields. Beam intensity, flatness, and symmetry status are monitored by the computer using analog signals from a thin Mylar window transmission ionization chamber. A given beam selection involves positioning six separate motorized, beam-modifying elements. They each must be properly aligned for accurate dose delivery. A number of interlocks are designed to detect any misalignment prior to abeamon condition. In this study the authors investigated the ability of the computer-controlled system to detect and respond to treatment beam dosimetry errors arising from misalignment, which
Linac operating parameters X-ray?
Energy, (Mv)
4
Electron?
Average beam current Filter in pA trans. (%)
Dose Rate (cGy/m/lm)
Energy (MeV)
Average beam current in nA
Scatter Foil (mil)
Dose rate, (cG~/m/lm)
3 Ta
500 500
6 10 15 18
200 100 70 50 30
45 35 30 25 18
2oob 400b 5OOC 5OOC 5OOC
6 9 12 16
100
25
20
10
5OOC
20
30
97 67 42
{ 1%
500
button
OlpA = A. = 1 0 - ~ n ~ b ~300 t pps, 3.5-ps pulse. CAt150 pps, 3.5-ps pulse.
500
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CHAPTER 9. DOSE MONITORING AND BEAM STABILIZATION
produced a nonflat beam, an asymmetric beam, an incorrect dose delivery, or some combination of these three factors. For conditions that may be encountered during normal treatment situations, symmetry, flatness, and dose delivery errors were detected and the radiation beam interrupted relatively quickly. However, for extremely artificial situationswith both the flattening filter and the backscatter shield (shutter) removed (see Fig. 8-2), the fault detection system failed. No fault or error message was displayed and the beam was not interrupted for conditions of excessive dose delivery. Although the experimental conditions described were extremely improbable, a QA check on the positional integrity of the flattening filter was instituted. The 3-mm A1 backscatter shield (shutter), overlying the incident surface of the monitor chamber for x-ray beams, appeared essential for the thin-walled Mylar ion chamber to respond appropriately to x-ray beam dosimetry errors.
TREATMENT BEAM STABILIZATION Effective and safe radiation therapy requires the performance characteristics of radiation treatment beams to remain constant during each individual treatment, as well as during a series of treatments extending over long periods of time. These characteristics can change with time for stationary field therapy. They can also vary with gantry angle and may change with gantry movement during arc or rotation (moving field) therapy. Their stability will depend on the design and construction of mechanical, electrical, electronic, and ancillary systems of the treatment unit. The operational requirements for treatment beams are best specified in the three IEC documents cited below. The publication IEC 976,21a a standard for medical accelerators, provides recommendations for the methods of test, and disclosure by manufacturers, of functional performance characteristics deemed necessary for radiotherapy. The publication IEC 977,21b a related technical report, provides guidelines for the functional performance characteristics for medical accelerators. These guidelines are recommendationsboth to manufacturers and users with respect to the performance of medical electron accelerators. They provide guidance to manufacturers on the needs of radiotherapists in respect to the performance of medical electron accelerators. They provide guidance to users wishing to check a manufacturer's declared performance characteristics, as well as for acceptance tests and periodic tests during the life of the equipment. The technical report IEC 977 includes a format for manufacturers to disclose their values of the functional performance characteristics together with a set of suggested values, which reflect the need for precision in radiotherapy and the knowledge of what is reliably achievable technically. An included rationale presents a concise view of the essential reasoning for many of the suggested values. It also includes a summary of test methods and conditions for acceptance and periodic tests and the recommended frequencies of
the latter. The International Electrotechnical Commission publication 601-2-1,21 is a general safety standard particularly relevant for the radiation safety aspects of patients and staff in the use of contemporary medical linacs. Of the many functional performance characteristics related to x-ray and electron radiotherapy, many are readily stabilized and need be measured only infrequently. Others, such as dose calibration, beam uniformity including flatness and symmetry, penetrative quality or beam energy, displacement of isocenter, congruence of radiation and light fields and dose rate are very important for accurate and safe dose delivery, are more apt to change, and require more frequent assessment. Calibration dose stability for x-ray and electron therapy is assessed by the variation of absorbed dose per monitor unit under specified conditions for both stationery and moving beams. It is first tested extensively during acceptance tests for high doses, as well as throughout a day, a week, and then periodically, and after relevant maintenance and repair work. Sirniliarly, beam uniformity, symmetry, flatness, isocenter displacement, congruence of radiation and light fields are assessed for their adherence within specified tolerance values on a scheduled basis as part of an ongoing QA program. The details of an individual QA program will depend on the type of equipment employed and the user's experience with it. Treatment beam stability often depends on electronic or electromechanical servo-feedback systems. Such systems monitor the treatment beam, or a related variable pertaining to the factor being servo-controlled.They generate an error signal which, when amplified, is used to operate a control element restoring the beam to within the required tolerance value for the particular factor. Servo-feedback systems are widely used in medical accelerators to control and stabilize important treatment beam factors. Several of the factors affecting beam stability, specifically those concerned with the incident angle and position of the electron beam on the x-ray target, or scatterer, are described in earlier sections of this chapter. For example, the electron beam is kept radially centered on the x-ray target and axis of the flattening filter by steering coils servocontrolled by signals from a sectored, ion chamber as described on pages 157-162. An advantage of such feedback systems is that tolerance limits, in this example in fractions of a millimeter displacement radially from the center, can be set by adjusting the gain of the amplifier incorporated in the feedback loop. Stabilization of electron beam energy, especially for bentbeam linacs, has been an important concern. A small change in beam energy, where such linacs employ 90" magnets, can produce a significant shift in beam direction, flatness, and symmetry in the treatment plane. An early approach to achieving beam energy stability was to servocontrol beam energy by varying the beam current by changing the electron gun emission. This alters the loading effect on the microwave fields in the accelerator structure, which results in an energy change of the accelerated beam. This approach was satisfactory for the peak beam currents needed for x-ray production, typically of
TREATMENT BEAM STABILIZATION
order 100 mA. For electron therapy, however, beam currents are smaller, typically of order 0.1 mA or less, and beam energy depends little on beam current and beam loading. Here, small changes in microwave source frequency differing from the frequency producing maximum output, can be used to vary beam energy. However, the beam energylsource frequency control characteristic is steep and the control method is not entirely satisfactory. Contemporary high energy linacs contain many treatment beam stabilization features and servo-control systems of varied complexity. These linacs usually employ 270" achromatic bend-magnets together with servoed beam-steering systems. Beam energy stabilization is provided by control of the pulse modulator supplying voltage to the microwave power source. The dose rate signal from the monitor ionization chamber is used to stabilize beam current and radiation output. Such linacs may provide both x-ray and electron treatment modalities with two or more x-ray energies and five or more electron energies. This versatility is achieved using the flexibility and stabilization provided by servo-feedback control systems. Beam symmetry may change with position of the gantry angle because of the influence of the earth's magnetic field over the electron path in the accelerator structure and the magnetic effect of the structural iron comprising the accelerator gantry, stand and base. Steel incorporated into the treatment room, as reinforcing rebar and thick primary barrier slabs, can exert a significant effect. Electrical arc welding of such steel can magnetize it and further perturb accelerator beams. The heavy steel magnetic shielding and magnetic fields of nearby MRI units can also be influential (see p 166). These effects will be most pronounced in the lower energy modes of machines employing long waveguides. To minimize these effects, the beam should gain a relativistic energy and the final energy over short distances of waveguide structure. Some systems provide adequate stability inherently without feedback. For example, the long-term voltage stability requirements of a power supply for an ion chamber operating well saturated need not be great. The performance of many subsystems and components can affect beam stability and experience has led to improved designs. Equipment modifications should be examined scrupulously by designers and engineers for their possible effect on beam stabilization. The findings noted and equipment employed in the study of treatment beam stability and anomalies have been extensively reported. Aird et al.2 found electron therapy output variations ranging from - 17 to +20 percent, depending on energy and field size, when the beam limiting aperture of the radiation head is covered with 2 mm of protective plastic. Even a 100-p Mylar cover, protecting the radiation head interior from dust, falling objects, and liquids (when upside down) affects changes of several percent depending on field size and energy. Dale11 cites an anomalous electron beam asymmetry problem. This problem was caused by the leads of the electron dose monitor chamber, which were separate from the x-ray chamber,sl sagging into the edge of the useful beam when the
165
gantry was rotated. The electron beam energy of linacs incorporating bend magnets can be inferred from the bend magnet current. The x-ray energy of low energy, straight-through linacs is usually monitored and feedback controlled by the ratio of peripheral to axial dose rate and checked from the ratio of ion chamber measurements placed at depths of 10 and 20 cm on the central axis. Gillen and Quillenl2 found that magnetron power in one such 4-MV linac, as assessed by magnetron current over its working range, caused a 10 percent change in output calibration and a 5-22 percent variation in flatness of large fields over a range of 1 to-10 cm depths. Padikal et a1.44 have described an apparatus for measuring symmetry during rotation using two off-axis detectors and an x-y recorder. Naylor and Williams39 describe an instrument designed for measuring small differences in dose rate at points in a radiation beam such as occur for small values of asymmetry. Loyd et al.31 studied the long-term variation in x-ray beam symmetry as a function of gantry angle for a Philips SL20 dual-energy, computer-controlled linac. They find day-to-day instabilities produce an overall variation in beam symmetry on the order of &2 percent suggesting that measurement of symmetry be incorporated in routine quality assurance procedures for this unit. Linacs employ water cooling to establish a stable operating temperature. Temperature is particularly critical for the ferrite materials of microwave circulators, a device that functions to prevent reflected microwave power from reaching and damaging the klystron or magnetron power source. A microprocessorbased temperature controller has been designed, which limits temperature excursions to less than 1°C for changes in operating conditions.37 Two 1970s studies focused on the x-ray and electron beam performance and stability of an early model radiotherapy linac incorporating a bend magnet and 360" gantry rotation.38.40 X-ray output constancy measurements exhibited 2 3 percent limits and flatness 25.5 percent limits for gantry angle effects during rotation.38 However, averaged over 4-week periods, a relevant interval for a course of radiation therapy, x-ray output constancy, and flatness were each within + 1 percent. Similar variations characterized electron therapy performance from 5 to 10 MeV.40 Here flatness limits were within ?4 percent, output constancy within 2 2 percent and energy constancy within 0.2 MeV over an 18-month period. By contrast, the stability of contemporary linacs is likely to be significantly improved from these values.4sa Radiation beam characteristics are measured, recorded, and incorporated into treatment planning under circumstances in which transient effects are eliminated. When treatments are given, the therapy unit is assumed to deliver radiation beams that match the data acquired during measurement. The ability of therapy units to reach their steady-state conditions of energy, output constancy, and field uniformity becomes a matter of increasing concern as the dose required for an individual procedure decreases. Often, such parameters as energy, field uniformity, and output constancy may vary significantly during the first second or two of operation and particularly during the first few MUs. The magnitude and implications of these per-
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CHAPTER 9. DOSE MONITORING AND BEAM STABILIZATION
turbations on patient treatment, port filming, and film dosimetry have been studied extensively for seven linacs and one cobalt teletherapy unit.4 Output and energy were evaluated using ion chambers placed on axis at dm,, 10 and 20 cm depths. Field uniformity (flamess and symmetry) measurements made use of a 30 ion chamber array connected to a microprocessorbased analyzer system providing both graphic and numerical display of data. Uniformity data were examined in both the radial (gun-target) and transverse directions. Most modem treatment units achieved steady-state x-ray beam energy (within a few tenths of a percent) in the first 2 MUs. One unit had a 6 percent initial energy variation but this was improved by reducing the pulse repetition frequency (PRF). Most units delivered a constant dose per monitor unit within a few percent at 10-cm depth after 8 MUs with smaller variations for increasing MUs. Straight-through linacs exhibit little or no field uniformity changes immediately after turn-on. Bent-beam linacs incorporate feedback stabilization circuits that need time to operate properly. The asymmetries appear to affect the radial (bending) plane and not the transverse plane. However, only one unit showed a large (23 percent) deviation in symmetry for the first 2 MUs as contrasted to a steady-state profile showing a deviation of less than 1 percent.4 The stabilization of arc therapy may involve either a constant or a variable output per degree of gantry arc rotation. A constant dose per degree can be achieved by varying either the output rate or the rotation speed in a compensatory manner. A variable dose per degree; for example, to compensate for changing thicknesses of overlying tissues, can be computercontrolled as a function of gantry angle of rotation. The time response of the system must be adequately fast to maintain beam stability during arc therapy. Electron dosimetry for arc therapy has been reviewed in the proceedings of a symposium edited by Paliwal45 (see also Refs. 34, 35, and 45 in Chap. 2 and p 41).
ELECTRICAL AND MAGNETIC INTERFERENCE Ion chamber electronics incorporate high impedance input circuits, which may be sensitive to electromagnetic interference. Kopecky and Purdy29 report interference with dose monitors from nearby microwave hyperthermia equipment at microwave field levels of 0.1 mWlcm2, a value considerably lower than an ANSI personnel safety standard of 10 mWlcm2 at 3000 MHz, the operating frequency of almost all linacs.3 Alternatively, van Rhoon et al.>3 found interference at 0.4 mWlcm2, from nearby hyperthermia equipment operating at 433 MHz with the beam energy determining electronics of another linac, but not with its dosemeter system. Aelectrostatically shielded room or a more distant location for hyperthermia equipment can provide solutions to these interference
problems. To reduce the likelihood of interference, all covers, doors, and panels on radiotherapy equipment should be kept in place, including keeping card-rack-housing drawers closed during routine use. Johnsen22 studied interference between MRI imaging systems and radiotherapy electron accelerators. Magnetic resonance imaging units can produce a fringe field of 3 G up to 15 m and 1 G at distances as great as 20 m. The effect of such magnetic fields on linacs is greatest in the regions of the gun where electrons have minimal magnetic rigidity. Experiments using a 3-G magnetic field resulted in less than 1/4 percent shift in symmetry at the isocenter of a Clinac 4 accelerator. The effect was greater for a Clinac 1800; 1*hpercent with servos disconnected but no noticeable shift with servos enabled. Linacs may also interfere with MRI units from fringe fields of their magnets as well as from the perturbing effect of their many tons of iron on the highly homogeneous magnetic fields of the MRI system. The fringing fields of a Clinac 2500 were less than 5 mG at 4 m and less than 0.1 mG at 15 m. A gantry, with approximately 6 tons of moving iron, creates a perturbing field of the order 0.7 mG at a 15-m distance. More than 100,000 new patients with heart disease have cardiac pacemakers provided each year in the United States. Lung and breast cancer together affect over 250,000 people per year, so that encountering a patient with both cancer and an implanted pacemaker in the same anatomic region is understandable. Pacemakers, usually implanted, serve as cardiac pulse generators. They are susceptible to external electromagnetic interference (EMI) as well as to damage from ionizing radiation. Electromagnetic interference may cause intermittant or transient malfunction while permanent malfunction might occur from accumulated ionizing radiation damage. Early pacemakers, which incorporated bipolar semiconductor devices, were found more sensitive to EM1 than to ionizing radiation based on a study of such devices in the field of betatrons and linacs.33 Aconclusion was that patients who have pacemakers implanted should not be treated with betatrons and linacs should be employed only with due caution since variations from one therapy unit to another can play a very important role in the type of interference exerted on the pacemaker. Contemporary pacemakers employ complementary metal oxide semiconductor (CMOS) for their integrated circuits because of the low power consumption, small size, reliability and The damage from ionizing multiprogrammable capa~ity.1~13*30 radiation in such devices is cumulative and failure may occur for doses of 10 Gy or even as low as 2 Gy.34 Hence, the pacemakers should not be irradiated; they should be placed outside the radiation field. The pacemaker manufacturer should be consulted if questions occur. The normal x-ray shielding of a facility provides sufficient attenuation to protect pacemaker wearers and patient monitoring equipment outside the treatment room. Physician monitoring of a pacemaker patient's pulse is recommended. The American Association of Physicists in Medicine (AAPM) has constituted a task group to report on relevant
REFERENCES
aspects of pacemakers in the radiotherapy environment. Preliminary recommended precautions in the management of radiation oncology patients with implanted pacemakers have been published34 and they follow below.
1. Have the patient's coronary and pacemaker status evaluated by a cardiologist prior to and soon after completion of therapy. Never treat a pacemaker implanted patient with a betatron. 2. 3. Always keep the pacemaker outside the machine collimated radiation beam, both during treatment and during machine portal filming. 4. Carefully observe the patient during the first therapy treatment to verify that no transient malfunctions are occurring. 5 . Estimate and record the absorbed dose to be received by the pacemaker (from scatter) before treatment. Locate the treatment fields so that the accumulated dose does not exceed approximately 2 Gy. 6. If treatment within these guidelines is not possible, have the pacemaker either temporarily or permanently moved. A detailed review of the pacemaker patient in the therapeutic and diagnostic device environment with extensive citations of the medical and electronic literature has been provided by Hardage et al.13 There is insufficient information available in the literature to allow detailed recommendations regarding the use of hyperthermia on patients with implanted pacemakers. Until such information is available, it is suggested that treatment be approached with extreme caution. Under no circumstances should the pacemaker or its leads be placed in the hyperthermia field or allowed to be heated above normal body temperature.
REFERENCES 1. Adamec R, JM Haeflinger, JP Killisch, J Niederer, P Jaquet: Damaging effect of therapeutic radiation on programmable pacemakers. Pace 5146-150, 1982. 2. Aird EGA, GD Lambert: Unforeseen variation of electron output with field size following modification of linear accelerator treatment head. Br J Radio158:1229-1230, 1985. 3. American National Standards Institute: Safety levels with respect to human exposure to radio frequency electromagnetic fields, 300kHz to 100 GHz. ANSI C95.1-1982 pp 1-24. 4. Barish RJ, RC Fleischman, YM Pipman: Teletherapy beam characteristics: "The first second. Med Phys 14557-561, 1987. 5. Boese HR, DV Cormack: Detection of a leak in a "sealed" monitor chamber. Med Phys 12: 377-378,1985. 6. Boux R: System for monitoring the position, intensity, uniformity and directivity of a beam of ionizing radiation. U.S. Patent 3,942,102, issued 1976. 7. Cheng P, H Kubo: Unexpectedly large dose rate dependent output from a linear accelerator. Med Phys 15: 766-767, 1988.
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8. Conere TJ: Some dosimetric discrepancies obtained using a guarded parallel-plate ion chamber with a high impedance electrometer in measurements involving pulsed and magnetically swept electron beam. Phys Med Biol31: 1157-1 160,1986. 9. Conere TJ, Boag JW: The collection efficiency of an ionization chamber in a pulsed and magnetically swept electron beam: Limits of validity of the two-voltage technique. Med Phys 11:465-468, 1984. 10. Cox RS, Heck RJ, P Fessenden, CJ Karzmark, DC Rust: Development of total skin electron therapy (TSET) at two energies on a Clinac 1800. Int J Radiat Oncol Biol Phys 18: 659-669,1990. 11. Dale RG: A possible cause of electron beam asymmetry in Siemen Mevatron linear accelerator. Int J Radiat Oncol Biol Phys 10: 1150, 1984. 12. Gillen M, R Quillan: Variation in output with fixed dose setting on a 4 MV linear accelerator. Med Phys 3: 107-108, 1976. 13. Hardage ML, JR Marbach, DW Winsor: The pacemaker patient in the therapeutic and diagnostic device environment. In Modem Cardiac Pacing, S. S. Barold, Ed, Futura Pub. Co. 1985; Chap 39, pp 857-873. 14. Holcomb L, DI Porat, K Robinson: Measuring position and current of an accelerated particle beam. Nuc Instr Meth 24399407,1963. 15. Hrejsa AF, J Soen, P Jankowiak: Temperature-pressure compensation for a linear accelerator electron beam dosemeter.Med Phys 12:26&261, 1985. 16. ICRU Report 21: Radiation dosimetry: Electrons with initial energies between 1 and 50 MeV, 1972, p 64. 17. ICRU Report 23: Measurement of absorbed dose in a phantom irradiated by a single beam of X or gamma rays, 1973,pp 30. 18. ICRU Report 24: Determination of absorbed dose in a patient irradiated by beams of X or gamma rays in radiotherapy procedures, 1976, p 67. 19.ICRU Report 34: The dosimetry of pulsed radiation, 1982, p 47. 20. ICRU Report 35: Radiation dosimetry: Electron beams with energies between 1 and 50 MeV, 1984, p 157. 21. International Electrotechnical Commission: Medical Electrical Equipment. Part 2: Particular requirements for the safety of medical electron accelerators in the range 1 MeV to 50 MeV. Publication 601-2-1, Section five: Radiation safety requirements. 47p., 1992. 21a. IEC Standard, IEC 976: Medical electron accelerators. Functional performance characteristics, 1989, p 147. 21b. IEC Technical Report, IEC 977: Medical electron accelerators. Guidelines for functional performance characteristics, 1989, p 129. 22. Johnsen SW: Interference between NMR imaging systems and radiotherapy electron accelerators. Med Phys 11:748, 1984. 23. Johnsen SW, P LaRiviere: Collection efficiency of thimble ionization chambers in intense pulsed x-ray and electron beams. Med. Phys. 14:491, 1987; and a private communication. 24. Karzmark CJ: Secondary emission monitor as a linear accelerator electron beam dose monitor. Rev Sci Instr 35: 1646-1652, 1964. 25. Karzmark CJ: Some aspects of radiation safety for electron acceleratorsused for both x-ray and electron therapy. Br JRadiol 40:697-703,1967. 26. Karzmark CJ: Procedural and operator error aspects of radiation accidents in radiotherapy. Int J Radiat Oncol Biol Phys 14: 1599-1602, 1987. 27. Kehoe TM, D Barnard: A testing jig for sealed beam monitors. Phys Med Biol32:649-652, 1987.
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28. Klevenhagen SC: Physics of electron beam therapy. Boston, Adam Hilger, 1985, p 204. 29. Kopecky WJ, JA Purdy: Radiofrequency interference with Clinac 20. Int J Radiat Oncol Biol Phys 6649, 1980. 30. Lewin AA, CF Serago, JG Schwade, AA Abitol, SC Margolis: Radiation induced failures of complementary metal oxide semiconductor containing pacemakers: A potentially lethal combination. Int J Radiat Oncol Biol Phys 10: 1967-1969, 1984. 31. Loyd MD, RG Lane, J Laxton, CH Chou, II Rosen: Long term variation in beam symmetry as a function of gantry angle for a computer-controlled linear accelerator. Med Phys 16:614-617, 1989. 32. Loyd MD, CH Chou, J Laxton, I Rosen, R Lane: Dose delivery error detection by a computer-controlledlinear accelerator.Med Phys 16: 137-139,1989. 33. Marbach JR, RT Meoz-Mendez, JK Huffman, PT Hudgins, PR Almond: The effects on cardiac pacemakers of ionizing radiation and electromagnetic interference from radio-therapy machines. Int J Radiat Oncol Biol Phys 4: 1055-1058, 1978. 34. Marbach JR: Recommended precautions in the management of radiation oncology patients with implanted cardiac pacemakers. ASTRO Newsletter VII-1,7-8, 1989. 35. McIntyre RD, CS Nunan: Linear accelerators systems having improved beam alignment and method of operation. U.S. Patent 3,838,284, issued 1974. 36. Meli JA, MS Weinhous: Collection efficiency of an ionization chamber in a pulsed swept beam: Chamber size effects. Phys Med Bio131:1139-1146, 1986. 37. Morris PB, MG Schmid: Accurate temperature controller for linear accelerators. Med Phys 11: 859-861, 1984. 38. Naylor GP, K Chiveralls: The stability of the x-ray beam from an 8 MV linear accelerator designed for radiotherapy. Br J Radio143: 414419, 1970. 39. Naylor GP, PC Williams: An instrument for measuring small difference in dose rate at points in a radiation beam. Phys Med Bio116: 525-528, 1971. 40. Naylor GP, PC Williams: Dose distribution and stability of radiotherapy electron beams from a linear accelerator. Br J Radiol 45: 603-609, 1972. 41. NCRP Report 69: Dosimetry of x-ray and gamma-ray beams for radiation therapy in the energy range 10 keV to 50 MeV, 1981, p 110. 42. O'Brien P, HB Michaels, JE Aldrich, JW Andrew: Characteris-
43.
44.
45.
45a.
46.
47.
48. 49. 50. 51.
52. 53.
54.
55.
tics of electron beams from a new 25 MeV linear accelerator. Med Phys 12:799-805,1985. O'Brien P, RB Barnett, HB Michaels, RA Siwek: Measurements of high intensity beams from medical linear accelerators. Med Phys 14:1067-1070,1987. Padikal TN, C Born, PL Robertson: The stability of teletherapy-beam symmetry with gantry angle. Radiol 139: 501-503,1981. Paliwal B, (Ed): Proceedings of the symposium on electron dosimetry and arc therapy. The Wisconsin Clinical Cancer Center, University of Wisconsin, Sept. 10-1 1,1981. New York, American Institute of Physics, 1982, p 373. Pamno PA, JA Purdy: The role of maintenance in quality assurance in Proceedings of a symposium on quality assurance of radiotherapy equipment. New York, Amer. Instit. of Physics, 157-168,1983. Pillar DG, MT Gillen, RW Kline, DF Grimm: A linear accelerator monitor unit totalizer. Med Phys 10:895-896, 1983. Sharma SC, DL Wilson, B Jose: Variation of output with atmospheric pressure and ambient temperature for Therac-20 linear accelerator. Med Phys 10: 712, 1983. Stieber VAW: Ionization chamber. U. S. Patent 4,131,799, 1978. Sutherland WH: Dose monitoring methods in medical linear accelerators.Br J Radiol 42: 864, 1969. Takata N, A Fukuda: Saturation characteristics of a vacuum chamber. Rad Prot Dos 8: 155-162,1984. Tauman L: The treatment head design for medical linear accelerators. IEEE Trans on Nuc Sci NS-28: 1893-1898, 1981. Tauman L: Dose monitor chamber for electron or x-ray radiation. U. S. Patent 4,427,890, 1984. Van Khoon GC, JA van de Poel, JA van der Heiden, HS Reinhold: Interference of 433 MHz microwaves with a megavoltage linear accelerator. Phys Med Biol 29:719-723, 1984. Weinhaus MS, JA Meli: Determining Pi,, the correction factor for recombination losses in an ionization chamber. Med Phys 115346-849, 1984. Weinhaus MS, JA Meli: Collection efficiency of an ionization chamber in a pulsed swept beam. Phys Med Biol31: 1147-1 155, 1986.
C H A P T E R
Accelerator Control and Safety Interlocking
Safe and accurate operation of the accelerator for patient treatment is implemented by the accelerator control and interlock system. The control console provides a central location for presetting, monitoring, and controlling operation of the accelerator for an individual therapy treatment. It may also incorporate selection of operational mode among, for example; morning check out, clinical, special procedures, physics, and service. It may contain digital displays including those for dose, beam-on time, as well as mechanical position readouts. Usually, it provides displays of dose rate and beam symmetry. The control console also provides: control of power to the accelerator, selection of radiation modality and energy, and other treatment parameters, as well as an indication of the status of various interlocks and subsystems. The console usually incorporates an illuminated panel display of interlock and subsystem status. Such a panel can convey 50 or more discrete status messages. Alternatively, an increasing number of computerized treatment units incorporate a video display terminal (VDT) to provide information concerning machine operation and the patient treatment prescription. Two key boards are often incorporated; one provides operational control of the treatment unit, the other provides patient data recall and modification. The console VDT provides an indication of normal operation or alternatively, information concerning an abnormal condition. Figure 10-1 shows a Clinac 1800 control console. It consists of a sloped, lower control panel and mounted above, a display panel of fault indicators, as well as digital and analog displays. The lower console control panel is divided into five (or six) groi~psof functional control and indicator modules. From left to right these are wedge and control power, arc therapy, x-ray therapy, beam control, and electron therapy. The extreme right module is reserved for control of optional modalities such as the high dose rate total skin electron therapy option. The display panel, mounted above the control panel, contains operational displays such as dose and dose rate together with interlock status lamps.
Interlocks, which are intimately concerned with the safety of personnel and with the proper operation of accelerators, form an integral part of the accelerator control system. Personnel interlocks are concerned with the safe and accurate delivery of the patient treatment prescription and the protection of radiotherapy staff and the general public. Machine interlocks are concerned with the safe operations of machines preventing damage to them. However, both personnel and machine interlocks perform their function by control of the treatment unit, that is, the machine. An interlock provides a means of preventing or terminating an operation unless certain predetermined conditions are fulfilled. Tert~zirlatingirradiatiorz requires stopping of irradiation without the possibility of starting without reselection of all operation conditions (that means return to the PREPARATORY STATE) (see p 173). Irzterruptirlg irradiation requires stopping of irradiation and movements with the possibility of continuing without reselecting operating conditions (that means return to the READY STATE) (see p 173). An abnormal or fault condi-
FIGURE 10-1 . Clinac 1800 control console. Lower, control panel. Upper, interlock display panel, showing fault status indicators, radiation modality indicaton: together with digital and analog displays.
10
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CHAPTER 10. SAFETY INTERLOCKING
tion, which terminates or interrupts a treatment, results in illumination of an annunciator having a key word or abbreviation informing the operator where the interlock chain is interrupted or which subsystem needs attention. Two examples of personnel interlock, fault messages are: ARC-lights when gantry stop angle is reached in arc therapy mode; TARG-lights when a discrepancy exists between the proper and actual target position (the target is extended into the electron beam in x-ray mode and retracted in electron mode). Two examples of machine interlock fault messages are VAC-lights when a significant loss of vacuum occurs in the accelerator waveguide; T/D-lights when a failure in the time delay circuitry occurs. Linac control and interlock systems increasingly employ integrated circuit (IC) electronics and digital circuitry in place of analog methods. These changes have improved reliability, decreased cost, improved performance, and reduced the dimensions of electronic systems, the latter often by a factor as large as 10. Replacement of discrete components within an IC is not routinely undertaken in place since the printed circuit (PC) boards, characteristic of the technology, can easily be removed and a duplicate board can be installed. This approach minimizes "down-time" and allows the defective board to be serviced under more favorable conditions. The digital systems interface more readily to computers than do analog systems. Interface electronics, which couple analog detectorsltransducers to digital devices, are increasing integrated into the overall electronics control system. The IEC publication 601-2-118 revision provides a wealth of technical recommendations for safe operations of radiotherapy linacs, particularly radiation safety of patients and staff. Greenel5 includes coverage of control and interlock systems. Swanson36 treats control and interlock considerations of radiotherapy linacs from the viewpoint of an overall safety program. Karzmark22 calls attention to the need for procedural safeguards to augment safety measures implemented by hardware or software.
COMPUTER CONTROL Computer control and monitoring of accelerator operation can provide many benefits. It can enhance safety by preventing bypass of interlocks and can capture, for later analysis, machine status, and operating conditions. Such information is valuable for servicing especially for infrequent intermittent problems, and can be interrogated remotely. It can provide the input data for expert system analysis benefiting from collective experience as described on page 254. The particular accelerator control system and operating program can be interrogated remotely. The prompt availability of these expert system problem analysis techniques can minimize downtime and increase machine availability. Computer control and monitoring can simplify acceleratorfunctioning, implement fast shut down and make it easier for the operators to carry out essential operations.
Using appropriate hardware, firmware and software, component tasks of an operation can be combined, choices emphasized and operations limited to safe, logical procedures. Hardware interlocks can be backed up with firmware. More sophisticated and complex treatment techniques become feasible under computer control of treatment units such as: Dynamic conformal therapy using multileaf collimators and/or independent jaws, universal or dynamic wedge and on-line portal imaging. Computer control of treatment units expands the benefits from computer integration of radiotherapy as described on pages 181-188. The computer based control system for a Philips SL 25 accelerator is illustrated in Figure 10-5. Computer network integration of radiotherapy is in an embryonic stage. Integration of diagnostic radiology will provide a solid basis of standards and techniques, from which methods for the more specialized application to radiation therapy systems can be developed. As examples of relevant conferences, a conference38 on "networks and image handling" was held July 3-4, 1986at theuniversity ofWales. Ajoint U. S./Scandinavianconference on computer aided radiotherapy was held in San Antonio, Texas, in 1988.20 A short course on computer management systems was presented at the 1988 World Congress on Medical Physics and Biological Engineering.30 Through the use of computers in treatment planning and in record and verify systems, as well as data base management functions, radiotherapy departments are already well advanced in the use of computers. It is a natural progression to want to tie together the various radiotherapy department functions by computer integration.
MINIATURIZATION Some idea of the degree of miniaturization present in accelerator control electronics may be gained from Figure 10-2; a PC board for a gantry position readout amplifier, where two levels of miniaturization are evident. Looking at the physical layout of the PC board, one sees eight analog integrated circuits (operational amplifiers and comparators) used for signal conditioning, as well as seven digital logic chips (gates and multivibrators) consolidated on a board measuring 6 X 4 in. Moreover, examining the inset equivalent integrated circuit at the level of a low power microscope, one finds that in each 3h -in. diameter operational amplifier "can," there is a sophisticated 20 active element device. Thus, miniaturization now permits techniques that would have been excluded from design consideration two decades ago on the grounds of reliability, physical size, heat dissipation, and cost. The diversity in application of IC electronics to linacs continues to increase. Their size continues to decrease, particularly, when viewed from the standpoint of functions performed. As the size and complexity of electronics systems grow, interfacing, interlocking, and system communication can become a significant problem. A low energy linac typically incorporates 20 PC cards, each associ-
SEMICONDUCTOR DEVICES AND ELECTRICAL INTERFERENCE
171
EQUIVALENT CIRCUIT
COMPARATORS (4)
OPERATIONAL AMPLIFIERS (4)
/
MULTIVIBRATORS (4) CHOKE,
RESISTORS (29)
28 PIN CARD EDGE CONNECTION GATE LOGIC (31
\
-
20 T U R N POTENTIOMETER (8)
TEST POINTS (5)
CERAMIC CAPACITORS (27)
POSITION READOUT AMPLIFIER BOARD FIGURE 10-2 . Printed circuit card for position readout amplifier. This 6 X 4 in. card contains tens of components. Some components, like the operational amplifier shown at the top, themselves contain tens of smaller conlponents in the form of integrated circuits (ICs).
ated with a specific task so as to simplify service. A high energy linac may incorporate 40 or more PC cards, several hundred individual IC units, as well as discrete components such as resistors and capacitors.
SEMICONDUCTOR DEVICES AND ELECTRICAL INTERFERENCE The potential for electromagnetic interference (EMI) around linacs is threefold: linac generated EMI, which interferes with
internal linac circuits such as control and dosimetry; linac generated EMI, which interferes with external electronic devices such as pacemakers and magnetic resonance imaging (MRI), and finally, EM1 sources external to the linac, which interfere with linac circuits. The internal sources are primarily the fast, high current pulses for magnetrons and klystrons, as well as the associated microwave bursts that may cause interference to other devices. The former are wide frequency spectrum sources with components up to about 50 MHz, and the latter, a single frequency of about 3000 MHz. Good circuit design, shielding, and grounding can minimize interference to linac circuits from both internal and external EM1 sources. In this context, all linac cabinet covers, doors, and panels should be kept in place or closed
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during routine linac operation. Optical couplers are frequently used for circuit isolation and simplification of grounding. Providing and maintaining a high quality,isolated ac power feed for the linac is recommended. Potential external EM1 sources should be adequately shielded and preferably located at a distance from the linac. Large scale integrated circuits, which involve many signal conductors and components in close proximity, are physically smaller than other devices, are more susceptibleto internal coupling, and great care is thereforeexercised in their design and layout. The dosimetry system itself may be susceptible to moderate levels of EM1 from external equipment such as microwave hyperthennia or diathermy equipment. Interference of this type may result in spurious counts on the integrated dose counters which could result in a dose misadministration to patients, an underdose or an overdose. Magnetic resonance imaging equipment is very sensitive to the presence of even small magnetic fields and nearby ferromagnetic material. Hence, such equipment and linacs should be located far from each other. Other aspects of electrical and magnetic interference are elaborated in Chap. 9, pages 166-167. The evolution of semiconductor devices has benefited linac control. Many safety and control decisions have a binary, digital nature; on and off, signal present or absent, and are often represented by numbers (1) and (O), respectively. Relays, a traditional binary device for control and interlocking, offer a high open circuit and a low closed circuit resistance compared to solid state devices, but have a higher failure rate. Relay coils are more likely to fail open circuit rather than closed circuit with solid state devices usually not exhibiting a preference for failure, either open or closed circuit. Although relays fail most often from a nonenergized,open-circuited coil, they also fail by contacts welding together. Contemporary miniature relays, moulded into a 14-pin chip mounting, consume about 200 mW of power. There has been a growing tendency to move away from relay control logic and to depend more on digital integrated circuits. Many machines employ hybrid control systems; a combinationof solid state logic and electromechanicalrelays. The early logic family, transistor-transistor logic (TTL), developed for computer use, was not suited for the high electrical noise levels found in linac environments. The margin between guaranteed output of one stage and acceptable input for the following stage, called noise immunity margin, was too small. A second family of logic was developed for high noise industrial environments called high noise immunity logic*(HiNIL), which operated at a higher voltage to improve the noise immunity margin. Figure 10-3 illustrates the voltage range of operation for 5 V TTL logic and 15 V HiNIL logic. The noise immunity range of HiNIL is approximately 10 times that of TTL in the high (1) state and is about 5 times larger in the low (0) state. High noise immunity logic is particularly effective in control applications where the ultrahigh speed of computer logic is generally not required and 15-V operation is commonly used in instrumentationand control equipment.Fast semiconductor devices have wide bandwidth with an attendant vulnerability to noise.
0.4 V Noise lrnrnunkv
.
FIGURE 10-3 Voltage reference diagram for 5-VTTL and 15-V HiNIL logic. Note the significantly increased noise immunity band for HiNIL at both the low (0) and high (1) state. Shading depicts input threshold range.
The noise environment carries great weight in the choice of a logic family for accelerator design. Not only are logic chips vulnerable to noise input, they produce noise that, in accumulation on interconnections, affects unrelated chips. Consequently, logic families have been developed, trading off various factors such as speed of response, power dissipation, and noise immunity. High noise immunity logic is costly, may be faster than needed, and dissipates significant heat. Where the noise environment does not demand HiNIL logic, it is now popular to use low power l T L or complementary metallic oxide semiconductor (CMOS) logic. Where possible, designers will consider using CMOS devices, which were originally developed for aerospace use. Complementarymetallic oxide semiconductor logic offers low power drain, permitting medium and large scale integration of function in one logic chip. While older CMOS logic permits 12 and 15-V operation and has a symmetry favoring low frequency noise immunity, the output is relatively high impedance compared with HiNIL and is vulnerable to electrostatic pickup. The newer CMOS has speeds approaching the older
INTERLOCK SYSTEM
HiNILlTTL logic families, requires only 5 V, but has the symmetrical design of CMOS. It is much less vulnerable to electrostatic interference and damage. Scanning one contemporary console with a cluster of about 170 logic chips, there is only one CMOS chip; and even with so much solid state electronics, there are 39 relays each packaged in a form approximating the area of a chip. The Clinac 18 uses HiNIL, with several exceptions, in 10 functional groups along with relay control. The logic groups are divided as follows: mode select and release, energy select, calibration and check, MU 1 display, MU 2 display, time control, system timing control, dose rate servocontrol, arc therapy, and interlock. The final interlock chain is based on relays. However, many of the coils of the relays are activated directly by integrated circuit logic.
ACCELERATOR OPERATIONAL STATES An accelerator, in being readied for and delivering a treatment, is carried through a hierarchical sequence of operational states: STAND-BY, PREPARATORY, READY, BEAM-ON, and COMPLETE or INTERLOCK. A medical electron accelerator, at any one time, functions in one of these defined operational states characterized by a varied and sequential readiness to provide a treatment beam. These states are accessed and usually displayed at the control console. STAND-BY is the state in which an equipment can be maintained for long periods, such as overnight and on weekends, and from which it is possible to move rapidly into operation. It is the state in which a working level of vacuum, temperature, and other parameters are maintained, but without the possibility to select the essential operating conditions. PREPARATORY is the state of equipment for setting essential operating conditions. These conditions are associated with a specific patient treatment and include setting the radiation type, modality, nominal energy, dose monitor units, and so. The setting of these conditions is precluded in the STANDBY state. READY is the state of equipment in which all conditions, such as the carrying out of confirming operations and any other satisfaction of interlocks, prevail so that the intended operation of such equipment can be initiated by a single action. An example of a confirming action is a second independent operation action of confirmation of the prior selection of the number of monitor units. The READY state is usually displayed on a lighted pushbutton switch or VDT display when the READY state is satisfied. The BEAM-ON state is the state of the equipment when delivering a radiation beam. The READY function is often combined with a COMPLETE function. This latter COMPLETE function, lights when the dosemeter has delivered the preset dose.
173
The INTERLOCK state indicates that an interlock fault is present preventing initiation of operation or that an interlock fault has occurred interrupting or terminating the exposure(see p 169). When interlock faults are present the equipment is moved to a lower hierarchical state requiring correction of the fault and may require reentry of the treatment prescription andlor other appropriate parameters. The particular operational state to which the equipment is moved will depend on the nature of the interrupt. An interruption of irradiation can, for example, be associated with opening the therapy room door. Door closure and confirmation of door closure at the console will restore the equipment to the READY state without the need to reselect operating conditions. A termination of irradiation fault interrupt moves the equipment to the PREPARATORY state, which requires resetting the specific patient treatment prescription operating conditions. Any dose delivered prior to such a fault interrupt is stored and accounted for in the reset operation unless the prescribed dose has been delivered. It will also be possible to terminate irradiation and machine movements at any time by terminating mains power to the equipment from EMERGENCY OFF switches, which may be located on the control panel, on the accelerator, on the sides of the treatment couch and equipment cabinets, and at strategic points in the treatment room. An EMERGENCY OFF switch interrupt disconnects all electrical power from the equipment so that a complete start-up procedure is required.
INTERLOCK SYSTEM The interlock system is designed to promote the safety of patient, staff, and public (personnel interlocks) and to protect equipment from damage during routine use (machine interlocks). The patient protection interlocks are designed to protect patients against all hazards associated with machine operations, largely radiation and mechanical hazards, and to ensure that the patient treatment prescription is accurately carried out. Protection is provided by prevention or termination of irradiation. The IEC publication 601-2-1'8 revision lists approximately 20 irradiation conditions that can initiate a patient interlock fault. Some specific interlocks for patient safety relate to: dose and dose rate monitoring, beam symmetry, beam energy, beam defining accessories, and computer control. Medical linacs employ extensive interlock and fail-safe design features for personnel safety. Their objective is to minimize danger to the patient and staff in the event of failure or malfunction of part of the system. Other interlocks enforce an operator control procedure sequence that minimizes the possibility of incorrect treatment of the patient due to operator error. Still others, called machine interlocks, protect the equipment from damage resulting from failure of a component or subsystem by shutting the machine off in the event of a malfunction. For example, the interlocks for the latter may be tied to mal-
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chain with a single illuminated indicator light. In addition, relay removal or failure turned on a fault light and inhibited operation. In such a system, a periodic manual lamp lighting check is used to establish the integrity of all lamp filaments. Interrelated circuits affected by the fault, yet requiring electrical isolation, were handled by auxiliary contacts on the activated fault relays. As the complexity of accelerator control grew, it was logical to move from the original 120 V ac relays to 24 V dc relays that were smaller and that required less power, interfaced with semiconductor sensors better, and operated on dc permitting use of semiconductor diodes as unilateral flow elements in the control switching. Moreover, in cases of trouble, the 120 V ac relays had the potential of releasing 120 V into the console with serious consequences for many solid state devices. Thus, it was a logical step in control circuit design to move to a diode switching matrix with semiconductor drivers such as transistors and silicon controlled rectifiers (see Fig. 10-4b). The advent of multimodal accelerators increased demands on the interlock system. During the same period, the reliability of semiconductor devices surpassed that of ordinary relays.
functions in the modulator, high voltage power supplies, the accelerator vacuum system, cooling water flow and temperature conditions, line voltages excursions, or other specific conditions. These equipment protection interlocks are incorporated into good design practices, but they are not required for patient or operator safety. Equipment cabinet access doors and protective covers are interlocked to protect staff from exposures to high voltages. Interlock testing is included in the QA program. Early medical linacs used an interlock system based on electromechanical relays with contacts in series to complete a "chain", as shown in Figure 10-4a. Each link in the chain was associated with a device or subsystem. Completing each link was indicated by turning on an associated and sequential indicator light on the console. By design, all fault indication lights were lighted during normal fault-free operation and with one or more lights extinguished to indicate a fault (s) when it occurred. Extinguishing a light identifies a fault and is also fail-safe against lamp filament failure. It was found more useful, however, to have the fault light panel normally darkened and to indicate the relative position of any break in the
lntlk B
lntlk C
Interlock
I
I
Aux. Function 1
I
Aux. Function 2
Sensor Sensor A B
Sensor C
Sensor D
Sensor
E
FIGURE 10-4 . Examples of accelerator control interlock systems: (a) relay ladder interlock and (6)diode matrix interlock.
INTERLOCK SYSTEM
This confluence led to solid state interlock systems in the form of computer logic components and diode switching matrices permitting great flexibility in design. The matrix can be driven with transistors able to follow rapid changes in the sensors or with SCRs (silicon controlled rectifiers) to create a latch requiring reset by the operator or auxiliary circuits. The diodes are highly reliable and yet inexpensive. Although the fail-safe nature of the design is somewhat compromised by transistors and SCRs, the rapidity of response permits inclusion of the matrix within a self-test cycle before operation. This self-test feature demonstrates readiness without adding appreciably to system delay. The interface to voltage comparators, to digital logic, to power-up states, or to relay contact closure is readily standardized. One standardized, general purpose hardware interface, CAMAC, as adapted for a Clinac 18, has been described.31 The example diode matrix interlock system sketched in Figure 10-4b provides both operational status (see p 173) and malfunction (fault) indication displays on the console concerning various circuits within the machine. A fault lights the appropriate lamp illuminating an abbreviated message identifying the fault on the console interlock display panel or a VDT (see Fig. 10-1). This system prevents beam initiation if an incorrect status is detected, or if a malfunction is detected by any of the sensors connected to the monitored circuits and to the vertical bus lines A-E of Figure 10-4b. In addition, this system causes immediate beam termination if the malfunction occurs after commencement of beam-on operation. The status signals from the sensors may denote an open switch on a protective cover or access door, or a malfunction associated with operator error or a linac operating parameter out of tolerance. A signal from a sensor causes the SCR (line A) or transistor (lines B-E) to conduct, thus sending an activating signal along its vertical bus. This signal is relayed by the cross connected diodes shown in Figure 10-4b to every horizontal interlock or auxiliary function circuit. The diodes serve to isolate faults from one another. Note that every sensor can interrupt the top interlock bus and its back-up (redundant) bus just below. The two interlock buses connect to series relays which, in turn, inhibit the beam, if any sensor sends an activating signal. Other sensor signals selectively activate the auxiliary function buses shown. These can inhibit beam-on and relate to such functions as dosimetry fault and usually involve complex conditional logic of specific subsystems. A sensor signal also activates an associated lamp shown at the top of each sensor bus. These lamps illuminate panel messages such as DOOR, when the door interlock is open, or VAC, when there is significant loss of vacuum in the accelerator waveguide. A total of 84 such messages can be activated in the Varian Clinac 1800 treatment unit as displayed in Figure 10-1. Many interlock functions are highly specialized as the following examples indicate. An excess dose rate interlock, defined at the normal treatment distance as more than twice the specification maximum dose rate, ensures that the dose rate does not significantly exceed the intended normal dose rate,
175
which could indicate a malfunction. This interlock is activated if too many monitor units are counted during a short time interval, or during a single pulse or a short pulse sequence, or during lo, or in a few degrees, of gantry rotation in the arc therapy mode. The excess-dose interlock is activated if the difference in monitor units of the dual dosimetry readouts exceed a predetermined value (e.g., 25 MU). The two dose integrators are functionally, and as far as practical, physically separate so that a failure in one system will not affect the other, and accidental intercommunication between the two systems cannot occur. Loss of delivered dose information, due to power failure, is guarded against by use of a back up mechanical counter or by auxiliary battery operation of the dosimetry monitor unit display (see Chap. 9 for more details concerning the dose monitoring system). The field symmetry interlock turns the beam off if the preset asymmetry limit is exceeded in either the radial or transverse planes of the radiation field. As noted in Chap. 9, the sectored ion chamber provides signals for beam steering designed to ensure operation within the preset asymmetry limits. An energy interlock can establish energy tolerance limits for both x-ray and electron beams using information from a sectored monitoring ionization chamber.13 Here the ion chamber has annular sectors arranged peripherally around a circular section centered on the beam axis. An angular distribution pattern of bremsstrahlung, originating in the x-ray target and modified by the flattening filter, will provide an energy-dependent signal based on the ratio of peak ion chamber currentlpeak target current, the ratio being a constant value for a given energy. Similarly, the angular distribution of electrons from a scattering foil will provide a constant value of this ratio, which is again a function of energy. Miller and van de Geijn26 describe a modification of the fault logic circuit for a Clinac 18 to accommodate large-field wedges. This circuit modification permits customer supplied large-field wedges to properly interact with the COLL fault interlock. As noted earlier, some control sequences at the console are designed to minimize the possibility of incorrect treatment of the patient due to operator error. These include: a reset control requiring important display readings (e.g., dose and time) be reset to zero or moved off the previous treatment values and then returned to the required values for the next treatment (even though the value may be the same as for the previous treatment), before a new treatment can begin. Several procedures require the operator to select among choices; for example, beam energy, electrons or x-rays, arc therapy, and wedged field. To ensure safety, many device activations require two separate and distinct actions called SELECT and CONFIRM. First, a SELECT operation is initiated, such as selecting and inserting the 30" wedge in the accessory mount in the treatment room. Then a CONFIRM operation must follow at the console, confirming that the 30" wedge is intended, before treatment can commence. The select and confirm operation can be extended to automation of treatment wherein the parameters selected for treating a patient must be confirmed by comparison with that patient's computerrecord before treatment can commence. The
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console may incorporate several keyed lock switches corresponding to the various treatment modes. The key serves a safety function. When inserted in a lock and turned, it identifies the treatment mode selected and allows the interlock chain to be completed. When the key is removed, the accelerator is rendered inoperable. An innovative use of integrated circuit logic occurs in the calibration and check cycle where a calibrate and functional check of all analog and digital circuits comprising the dosimetry system are carried out prior to treatment from each field. After the operator selects one of four different modes of clinical operation (ARC CW, ARC CCW, FIXED X-RAY, E-BEAM), a choice of energy is made, whereupon, an automatic calibration and check sequence is initiated. After a logic reset, a 6-s calibration sequence occurs with both dose counters counting to 100-109 MU, the timer counting to 0.1 minute, and the five excess dose rate faults are all activated and checked. The excess dose rate fault modes relate to exceeding specific preset limits: MUIDEGREE, MU IITIME, MU 2/TJME, MU IPULSE, MU 2Pulse. A small electrical current is applied to the ion chamber plates of each dosemeter for this latter check. The dosemeter calibrate interlock check will be completed only for a consistent set of integrated current readings from the two dosemeters. After this calibration, the logic is again reset and both dose counters and the time counter are fed 360-Hz signals and allowed to count up to the dose indicated in the thumbwheel dose settings as the check sequence. Then, the third reset occurs, whereupon the arc therapy logic and both dose integrators are enabled. Before treatment proceeds, the logic demands a verification of the chosen energy by requiring the operator to confirm by reselection of energy.
PROTECTION AGAINST EXTREME DOSE Multimodality treatment units that provide both x-ray and electron beams, potentially present a particular extreme dose rate hazard in that the high electron beam currents needed for x-ray production, especially 4-6 MV, can result in extremely high dose rates if not intercepted by the x-ray target or flattening filter. Although the higher energy x-ray modalities entail lower beam currents (but still higher than comparable electron therapy energies) they too present a hazard because of less scatter of the electron beam in the foil, window, and ion chamber. An extreme dose rate hazard is defined as more than 10 times the vendor specified maximum dose rate at normal treatment distance. For example, 100 pA of average beam current may be employed for 6-MV x-rays but only 0.1 pA of beam current for an equal dose rate of electrons at 4 Gylmin at 1 m (see Table 9-1). An abnormal fault condition could involve machine operation in electron therapy mode but with 100 pA beam current typically used for 6 MV x-ray therapy. Under this fault condition electron dose rates of 4000 Gylmin are possible
and constitute an extreme hazard for patients. Such an electron beam, particularly if unscanned or unscattered, can collapse to a small diameter less than the central electrode diameter of the monitor ionization chamber. In this circumstance its response would erroneously indicate a much lower dose because of this geometric anomaly as well as a likely lowered ion collection efficiency. Presently, IEC recommendations18 state that if treatment units can deliver under any fault conditions a dose rate at normal treatment distances more than 10 times the vendor specified maximum dose rate, then the unit shall incorporate a beam monitoring device, which is independent of the dose rate monitoring system and located on the patient side of the x-ray target and beam distributing system. This device shall prevent the absorbed dose rate from exceeding twice the vendor specified maximum dose rate at normal treatment distance and will limit this dose to less than 4 Gy anywhere in the radiation field. The beam monitoring device must respond to failure of a scattering foil or lack of movement of a scanned beam. To accomplish these objectives, the device will need to monitor the beam on a pulse-by-pulse basis and terminate the exposure within a time on the order of lms. An additional IEC requirement for extreme dose rate interlock involves either: termination or prevention of irradiation when an interlock component fails; or testing of interlocks between irradiations; or redundant or continuously monitored diverse interlocks. An analysis of IEC publication 601-2-118 lists approximately 30 interlock fault conditions that can result in prevention or termination of irradiation. Of these, roughly 20 can give rise to the extreme dose hazard identified herein. The radiotherapy beam characteristics of one multimodality 25-MeV scanned electron beam linac have been described by O'Brien et al.29 For this linac in the extreme dose rate fault condition with the abnormal unscanned electron beam intercepting only the monitor ionization chamber, an extreme dose of between 1 and 2 Gy per pulse was produced at normal treatment distance at or near a depth of Dm,,More than 100 pulses were emitted before the interlock interrupted the beam. This scanned beam linac is computer controlled and the abnormalextreme dose fault involved software and system response time. Loyd et al.25 assessed the dose monitor errors in the Philips SL25 computer-controlled linac (see Fig. 10-5). They found that for conditions that may be encountered during normal treatment situations that the symmetry, flatness, and dose delivery errors were detected and the radiation beam was interrupted relatively quickly. However, for extremely artificial experimental situations with both the flattening filter and the backscatter shield removed the fault detection system failed (see Fig. 8-2). The current draft IEC safety standard18 requires computer hardware and computer programs that control the system to be capable of fulfilling all safety requirements of the standard under all conditions, including transient or permanent failure of the computer or related interface. Redundancy of critical safety interlocks is essential with different physical principles being employed in the sensor portion and different
177
CONTROL CONSOLE
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. Accelerator control system for an SL25 linac. (Courtesy of Philips Medical
logic being employed in the comparator and actuator portions of each such interlock. It should be noted that many conventional radiotherapy calibration dosemeters are often inadequate for the extreme dose rate measurements. One commonly used 0.5-cm3 ionization chamber exhibited collection efficiencies of only 10 to 15 percent under these conditions.29 Additional aspects of beam monitoring and control of multimodality treatment units under normal and abnormal conditions18 are presented on pages 162-164.
CONTROL CONSOLE The control console is designed to simplify ongoing operation of the treatment unit. It is the central location for presetting, monitoring, and controlling operation of the treatment unit. It provides for selection of beam type and energy. It is equipped with lighted indicators that show the status of equipment interlocks. It usually provides displays for dose rate, integrated dose for both primary and secondary dosimetry channels, treatment time, gantry angle, arc therapy and rotation direction (CW or CCW), and operating mode (either x-ray or electron; and fixed, full-field or arc therapy). In the fixed and full-field modes, the gantry remains stationary. The full-field exposure is used to provide anatomical references by superposing a maximum collimator opening (full-field) exposure on a treatment port film radiograph. In the arc therapy mode, the gantry rotates
during treatment. In multimodality units, additional displays and controls pertain to radiation type (x rays or electrons) and beam energy. The console logic monitors and displays the status of various circuits within the machine to provide both operational status and malfunction (interlocWfault) indication. Beam activation is prevented if an incorrect status is detected or if a fault is detected in one of the monitoring circuits. Immediate beam termination ensues if the malfunction occurs after comrnencement of "beam-on" operation. An interlock and status matrix assembly provides monitoring and fault indication from a sequence of labeled fault indicator lamps on the control console as described on pages 173-176. These lamps are activated by fault detectors placed in various locations throughout the machine including the console and various PC boards. Stevens et al.34 describe a digital satellite addition for the Clinac 18, which displays the gantry and collimator positions at the console, a feature particularly valuable in setting the back-up x-rays collimator jaws 5 cm larger than the electron applicator selected for use. There is increasing use of computers and microprocessors in linac control systems. In the late 1970s, the AECL Therac 6 treatment unit began to employ a Digital Equipment Corporation (DEC) PDP-11/05 computer to monitor and control machine operation.14 Its video display unit could display a clinical treatment format of each patient's prescription or a maintenance format of the current value of 64 linac operational parameters. This latter display can be useful in rapidly diagnosing or anticipating component failures, as well as in optimizing performance. Computers can also be used to preset the treat-
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ment positional variables on a daily basis to the prescription values. They can also provide the basis for dynamic treatment via programmed dose, collimator jaw, and gantry motions, as well as coordinated movement of the patient couch with respect to the radiation beam during treatment (see Chap. 2, p 41-43). A computer-controlled therapy accelerator must be extensively tested to ensure that the specific hardwarelsoftware system involved enhances safe operation and does not permit unsafe practices or anomalies. An extensive testing procedure for one prototype accelerator,a Varian 2100C, has been carried 0ut.39 The configuration of the 2100C at any time will depend on the operational mode (e.g., morning checkout, clinical, physics, and service), the radiation mode (x ray or electron and energy) and the operational state (stand-by, ready, beam on, etc.) (see p 173). The selected operational mode, radiation mode, and operational state properly constrains how the accelerator can be operated. In clinical mode, the physicslservice alphanumeric keyboard is rendered inoperative and the operator's keyboard will respond only to an appropriate and restricted set of commands associated with the operational state as indicated by an on-screen display. The acceptance testing procedure included mechanical systems, radiation parameters, and manual safety systems, but the emphasis of this report was on the computer control system (communications integrity, state integrity, etc.) and the interlock systems (some 50 electrical, mechanical and/or computer controlled interlocks). Specific tests were related to adherence to design specificationsfor normal operations. Some tests constituted attempts to produce abnormal operation and others to assess computer/interlock interactions. An appendix contained in the above noted report,39 provides details of procedures for testing interlocks to assure safe operation for patient and machine. The procedures relate to patient dose monitoring and other interlocks together with the computer control system. All safety related interlocks were tested in the clinical (patient treatment) mode. In clinical mode, interlocks cannot be overridden. Figure 10-5 is a functional block diagram of the accelerator control system for a Philips SL25 treatment unit. The SL25 provides two selectable x-ray energies between 6 and 25 MV, together with nine electron energy beams ranging in energy from 4 to 22 MeV. The dashed blocks included in Figure 10-5 constitute the hardware associated with the record and verify option. Much of the hardware, including three 16-bitmicroprocessors, common memory, and storage devices, is housed in control cabinet "A" and interconnects with the accelerator and console. The operator console is comprised of a video display unit and associated keyboard, a TV room monitor, and a keypad that directly controls the accelerator functions such as, start, interrupt, and stop. Processors 1 and 2 control data communication with the accelerator and accelerator functions. Processor 3 controls access to the Winchester disk and other peripherals. A35 Mbyte Winchester disk forms the main storage for all programs, patient data, and machine data. The 1 Mbyte floppy disk is used for loading the operational and test software during the instal-
lation and maintenance of the accelerator, as well as for the transfer of treatment prescription data. The keyboard is used to enter a treatment prescription for every new patient that includes: patient identification and nonmechanical parameters (e.g., radiation type, energy, and monitor units). Mechanical parameters (e.g., field size, gantry, and collimator angles) can be entered manually or transferred from the linac when the patient is correctly set up. This treatment prescription is then stored on the Winchester disk for future sessions, or'deleted if it is not to be used again. The current patient prescription is displayed on the TV monitor in the treatment room, as well as on the TV monitor at the console. Patient treatment prescription verification with customized tolerances together with updating of the patient record, both on the Winchester disk and a hard copy record, are provided as described on page 180. Computer based accelerator control systems facilitate the implementation of additional features that may include: Computer-assisted setup of patients, remote-controlled, dual-exposure check radiographs (6 MV) for port films, secondary access of treatment data from a remote terminal, computer-assisted service for maintenance including automatic self-logging of machine data, self-diagnostic facility via phone line communication using a modem, arc (moving beam) and dynamic therapy, and linkage with an external computer up to 1000-m distant. A current trend in treatment equipment design is to incorporate ever larger numbers of microprocessors, each with an assigned function. These function-assigned microprocessors can operate under the jurisdiction of a master computer. Protection against the extreme dose hazard of dual-modality (x-ray and electron) linacs, which assumes a larger role in computer-controlledunits, is described on pages 176177.
MOTION CONTROL SYSTEM The Clinac 1800 motion control system is associated with positioning the equipment for patient treatment; motions of the gantry, collimator, patient support assembly (PSA), or couch and beam stopper. In addition, it includes motions of subsystems that are essential for operation of the accelerator to enable the correct radiation treatment modality. These latter functions are depicted in Figure 10-6, a motion system block diagram. The associated motions are implemented by drive motors, stepper motors, or pneumatic actuators. As shown in Figure 10-6, they affect the waveguide shorting tee (Chap. 5, p 98) the x-ray target, as well as carousel and mirror positions. These motions are activated when the beam energy and mode, (x rays or electrons), are selected. Associated logic and interlock blocks are also shown. The treatment of patients involves motions of the gantry, collimator, patient support assembly, and beamstopper. These
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MOTION CONTROL SYSTEM
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. Motion system block diagram for accelerator operation.
motions are controlled by nine drive motors shown on the right of Figure 10-7; the motor system block diagram. The gantry can rotate through a range of 185" and has a continuously variable rotation speed from approximately 0 to 1 rpm. The collimator assembly rotates around the central axis of the beam through an angle of -+90°. The collimator field size is continuously variable from 0 X 0 cm to 35 X 35 cm at the isocenter at 1-m distance. The patient support assembly rotates 95" around a vertical axis, which passes through the isocenter. The couch translates laterally 5 2 5 cm, longitudinally a total of 110 cm, and provides a couch table top vertical range from 82 to 135 cm above the floor (see chap. 12, pages 201-203 and Figure 12-1). The beamstopper moves into and out of position within 60 s. The various position motors found in the stand, PSA and gantry are controlled by circuitry located throughout the system. However, these control functions are directed to the motor control subchassis (see Figure 10-7) where the signals interface and control the power to the individual drive motors. The motors are dc operated for ease of control. Some motors have permanent fields and others are energized from dc field coils. The direction of rotation of the motors is altered by reversing the polarity of a sinusoidal dc through the armature of the motor. The speed of rotation is varied by the timing of a silicon control rectifier SCR trigger pulse, which determines the time during the half-cycles at which the partial sinusoid is applied. The various motorized motions concerned with treatment beam positioning are primarily controlled from a hand-held
+
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pendant but some motion controls are duplicated at the radiation head, the control console, or the patient support assembly. The pendant has four thumb-wheel potentiometers, whose displacement from their center resting (deactivated) position determines the direction and speed of the motions they control. In addition, there are three operational slide switches. One switch controls pendant mode, directing control to the couch or gantry. In couch mode, the thumbwheels control couch lateral, longitudinal, vertical, and angular motions. In gantry mode, the thumbwheels control gantry rotation as well as collimator rotation and field size. The other two slide switches control the radiation field and room lights. The pendant also contains a dead-man switch, which must be closed for any motion to occur. The PSA has duplicate controls on the couch for lateral and longitudinal motions. The motor drive for these latter motions may be unlocked for manual positioning of the couch. The collimator assembly contains motor controls for adjusting field size and collimator rotation angle together with mechanical readonts. In addition, a large circular panel in the yoke of the gantry displays both gantry and collimator position data in digital form. Control of a full-field mode of the collimator jaws for portal filming is located at the control console. An alternate pendant design employs membrane type linear potentiometers. The latter takes the form of a flat, normally open (deactivated) device and is configured as a standard three-wire potentiometer. It is activated by pressing the membranes and sliding one's thumb or finger laterally across the pendant at the location of the potentionmeters.
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The analog output from the potentiometer signal may be digitized or used directly.
RECORD AND VERIFY SYSTEM Small computers and microprocessors can be used to monitor, control, and place limits on linac operational parameters, as well as store and manipulate information pertaining to patient treatment. They can enter a record of an individual patient's treatment prescription in computer memory, verify the correctness of the daily sequence of treatments, as well as record the daily treatment doses and their cumulative sums at specified anatomical sites. 638321224,27,28,32,33 Table 10-1 lists 14 parameters monitored in one record and verify system.33 Under each parameter are given its range of adjustment and resolution. The resolution of analog channels 1-8 refers to the potentiometer linearity and does not include mechanical error. Unless the patient can be reproducibly repositioned on the treatment couch on a daily basis, a difficult and exacting requirement, it is not useful to include such positioning parameters in the system. A further extension of record and verify systems is dynamic therapy under computer control as described in Chap. 2, pages 41-43.
Although the expression "record and verify" (R&V) is in widespread use, it is a misnomer. The sequence "verify and r e c o r d correctly describes the actual order of first verifying the treatment prescription and individual daily doses prior to recording them and their contributions to cumulative sums. A number of centers have investigated such record and verify systems, and they are now being offered as an option by manufacturers. Their ability to identify errors in machine setting of the treatment prescription, errors that can be corrected prior to daily treatment, could be a significant benefit. The types of errors found, their magnitude and frequency of occurrence have varied at different centers and may be unique to a particular center or a specific type of ma~hine.12.21~32~33 An early study of the automation of radiotherapy treatment machines is that of Enviro-med.9 Their study looked at 4 levels of automation and their effects on patient care, operational efficiency, and costs. Early studies of the error rate in setting treatment machine and couch parameters were carried out by Kartha et al.21 who found timer or monitor unit setting responsible for the largest component. The error rate will depend on the preset tolerance values of treatment machine and couch parameters, which allow treatment to begin and with small errors predominating. The record and verify systems offered by manufacturers differ significantly in their features and are in general incompatible with each other. Mohan et a1.27,28
COMPUTER INTEGRATION OF RADIOTHERAPY
TABLE 10-1 ---
. Monitored parameters
-
Parameters
Gantry angle Upper collimator jaws Lower collimator jaws Collimator angle Vertical couch position Longitudinal couch position Transverse couch position Couch angle Dose per degree Dose Time Arc stop angle Wedge Shadow tray
Range of adjustment
Resolution
0"-360" 0-32 cm 0-32 cm -C 90" 2 2 cm 0-141 cm 225 cm -C90° 0.5-5 radP 0-999 rads 0-9.9 min 0"-360" 0-7 IN or OUT
0.2% 0.2% 2% quantization levels 1cGy 0.1 min 1 O
1
describe a standardized interface, which facilitates treatment monitoring of diverse accelerators on a single central computer. Large errors in field size, timer, or monitor setting, which could significantly affect overall dose prescription, appear to occur infrequently.12.21If so, they may have only a small effect on the outcome when averaged over a course of therapy. Therefore, the equipment and effort involved to identify and correct such errors may not be justified, except in special cases, for example, spinal blocks and lung blocks. However, with treatment becoming increasingly complex and with the wider use of dual x-ray energy and dual modality techniques, a record and verify system, may become an essential ingredient of the QA program of a treatment center. Alternatively, one may incorporate the record and verify function into a larger computer system which includes accelerator control and monitoring, together with extensive record keeping. The rapid advances in computer hardware and software have benefited all radiotherapy computer applications. Commercial record and verify systems are becoming more sophisticated and may include extensive data management capability. They often provide color window displays, popup menus, data entry from simulators, remote review and editing workstations, and incorporation in local area networks.
PATIENT RECORD KEEPING A significant extension of the record and verify feature is to replace entirely the manually kept daily therapy record and concomitantly reduce the record keeping burden for technologists. Relieving treatment technologists from the responsibility
181
of maintaining the daily therapy record of treatment appears a fruitful application, but there is limited experience. Certainly, technologists, if freed of this responsibility, which includes arithmetic aspects, would have more time for observing the patient and for patient concerns. However, it is essential that the patient's current record, which is stored in the computer, be available at all times and particularly, when the computer is "down." At least one commercial system appears to have solved this problem by employing a formatted and printed daily therapy record form incorporating an integral magnetic strip record and a terminal specifically designed to update it.8 For each treatment session or other entry, the individual patient's record form is inserted into the terminal and automatically indexed to the appropriate printing position for updating it, the magnetic strip, as well as the computer record. Normally, the patient's daily therapy record would be quickly viewed on a video display unit. However, in the event of computer malfunction, the current printed record is still readily available. Another difficulty in such computer based systems is designing software programs that can handle all, or almost all, of the routine entries, as well as contingencies that can modify the patient's daily record during a course of therapy. Even a modest need for manual intervention may render a record and verify system more of a burden than a benefit.12 Software must also be readily modifiable to accommodate changes in administrative procedure of treatment technique. Patient record keeping involves an essential core of information but it, together with any extensions, depends a great deal on personal preference of format, conventions and style.
COMPUTER INTEGRATION OF RADIOTHERAPY The term computer integrated radiotherapy can be defined as the digital transfer of information among individual functions (applications) under computer management. Individual applications require data input, often from another application, and generate data output, which is often filed for later access or transferred to another application. As opposed to manual methods, integrated therapy can facilitate the transfer of image and other data with minimum distortion and can facilitate file management and access. In order for different individual applications to communicate with each other, standardization of transfer data format is essential, with interfaces to adapt the communication system, to individual application equipment.1.2J7 Horiil6 describes one example of such a format, the (ACR-NEMA) American College of Radiology-National Electrical Manufacturers Association standard for transfer of diagnostic images. Record and verify systems have been offered by most manufacturers of radiotherapy machines. Such systems provide for storage and retrieval of patient and treatment informa-
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tion, prohibit treatment outside user-defined tolerances on machine settings, and provide for automatic setup of selected machine positions and functions. l'j~picallyin the past, each treatment machine had its own individualizedrecord and verify system. Mohan et a1.28 describes a centralized record and verify system, which serves four treatment machines of three different manufacturers. Intelligent interfaces were developed to take input from the machines in their different formats and communicate it digitally to a host computer in a standard format. The host computer inhibits the beam until the parameters set on the treatment machine agree with the prescribed parameters for treatment within specified tolerances. Patient demographic information, treatment prescriptions, individual treatments, and verification failure out of tolerance data are recorded and used to generate reports that are available for display or printout. Record and verify systems are available that can serve more than one treatment machine, provided they are by the same manufacturer. Such systems employ a computer and hard disk for verification, recording, and database management. Manufacturers (e.g., see Ref 4.) are working on the develop ment of systems of broader capability, in order to provide digital transfer of patient treatment data among several functions in the radiotherapy department, such as patient contour sources, treatment planning system, simulator, treatment machine, physician work station for image display and manipulation, and terminals for administrative tasks and office automation activities. Table 10-2 lists a rationale for computerization and some specific applications of computers in radiotherapy. A variety of local area network protocols are available for communication among functions. One example is Ethemet.10 It employs a coaxial cable to which all hosts connect through interface boards. A machine wishing to communicate with another waits a randomly chosen time (repetitively, if necessary) for the cable to be free, then broadcasts its message on the cable. The modulation rate is 10 Mbitsls, providing data rates of 0.5 to 5 Mbitsls. Cable length is limited to 500 m. Communication is in "baseband" (direct pulsing). Some other systems operate in "broadband" (modulation of a carrier, like radio transmission) permitting higher data rates, but at greater expense for interfaces. The future development of fiber optics networks will facilitate high speed transfer of high resolution images. Diagnostic information from many sources is used for radiotherapy treatment planning. Much of this diagnostic information is available in digital form, such as images from x-ray computerized tomography (CT) scanners, MRI, digital radiography (DR) systems, nuclear cameras and scanners, and ultrasound scanners. It is essential that such image data be provided to the radiotherapy treatment planning computer without distortion. Adigital imaging and communicationsstandard' developed by ACR and NEMA will: 1. Promote communication of digital image information regardless of source format or device manufacturer,
2.
Facilitate the development and expansion of picture archiving and communication system that can also interface with other systems of hospital information, 3. Allow the creation of diagnostic information data bases that can be interrogated by a wide variety of devices distributed geographically. A network may or may not be involved. The ACR-NEMA standard specifies the hardware interface, certain software commands, and a set of data formats for communication across an interface between an imaging equipment and a network interface unit or another imaging equipment. It is not an overall picture archival and communication system (PACS) or work station specification, nor a network standard. The ACR-NEMA standardl.2.16 defines a point-to-point hardware connection with protocol and data structure such that two differing devices (e.g., imaging device, workstation, or laser film writer) will be able to communicate with each other, exchanging images and associated data. On each side of an interface, the same layers are performing the same functions, so the programmer views each layer as communicating with its equivalent. The communication actually occurs in a vertical direction in the layers and the true connection is only at the TABLE 10-2
Computerization-rationale-specific applications
Rationale I. Improve quality through improved treatment planning and verification of set-up and dose delivered to the patient. 2. Perform tasks not possible manually (e.g., dynamic control.) 3. Increase the volume and type of data available by generating data bases, for statistical analysis and control as well as patient management such as scheduling and billing. 4. Create patient files with patient data, diagnosis, course plan, and summary of treatment given. 5. Reduce the time needed for routine (but critical) data recording and checking. 6. Include radiotherapy accelerator diagnostic information and facilitate thorough rapid quality assurance checks. 7. Manufacturing upgrades become software rather than hardware changes. Some spectfic applications
Record and verify. (Enter patient data and course plan, compare set-up parameters with plan, record results.) Treatment planning. Data management. Radiotherapy accelerator control system. Dynamic (conformation) therapy (advanced computerized treatment planning; computer control of accelerator). Automatic setup of accelerator (with record and verify system). Automated dosimetry measurements, storage of data in computer, and computer presentation of data through graphs or printouts. Dosimetry control through use of look-up tables containing calibration and output factors.
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COMPUTER INTEGRATION OF RADIOTHERAPY
physical layer. The message format itself consists of data elements collected into groups. Each data element (e.g., patient name) consists of two numeric fields forming a key. This is followed by a length field and the element value. Once the message is assembled, it is sent down to the session layer. The transportlnetworklayer takes the message with the information from the session layer and fragments the message packets. In the data link layer each packet is enclosed by a frame descriptor word and a frame check sequence to form a frame. The frames are sent to the physical layer, which handles transmission across the interface. The physical interface consists of 16 asynchronous parallel lines with control signals. Figure 10-8 shows the data flow through the ACR-NEMA interface. Figure 10-9 shows a schematic of a radiology storage transfer analysis and reporting (Q-RSTAR) system employing picture archiving and communicationssystem (PACS) technology in a department of radiology?' A dual local area network (LAN) is employed, comprising Ethernet coax for command and control information at data rates to 0.2 Mbytesls and an optical LAN capable of sending high resolution images in a timely fashion at up to 100 Mbits (e.g., 12 Mbytes 8 bits deep). Figure 10-10 shows Ethernet system components.3 In radiotherapy departments the volume of images is much
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less than in a radiology department so an optical LAN is not essential for image transfer, such as from a simulator and electronic portal imager to a workstation. However, compatibility with the radiology department in transfer of CT, MRI, and so on images for treatment planning may encourage use of optical LAN in the radiotherapy department; especially for the quantity of images needed for conformal therapy and 3-D treatment planning. In addition to the ACR-NEMA effort, which was initiated primarily to serve the needs of diagnostic radiology, the IECl7 has started work on standardization of data exchange for systems in radiotherapy. This IEC effort is intended only to standardize the terms used in transmitting information and their mnemonic abbreviations. Dahlins has reported the initiation of a Nordic program among medical centers to develop an integrated information system in radiotherapy, called CART (computer assisted radiation therapy) (see Figs. 10-11 and 10-12). A logical information flow is defined, which corresponds to the standard sequence of radiotherapy procedures: 1. Image processing: Localize tumor distributions. Define target volume.
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50 R Coaxial Cable
Taprrransceiver Assembly
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5.
Oncological data base: Transfer to hospital storage of all essential data to permit long-term follow-up and statistical evaluation of different treatment modalities.
The sophistication of treatment planning will develop over time and any future integrated information system should be planned to anticipate the following future treatment planning capabilities:
Interface
-
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. Ethernet system components block sketch
Define patient outline and organ contours. Define physical properties of treatment beam. Subsequent follow-up studies for tumor and organ response. 2. Treatment modeling: Acquisition of dosimetric beam data for source to be used. Obtaining of data for making patient individual beam modifiers. 3. Treatment verification: Treatment unit set-up parameters. Patient specific devices, including beam modifiers. Patient positioning. Absorbed dose to the patient. Treatment unit electrical and dosimetric performance. Documentation of treatment prescription. Optional automatic set up of the treatment unit. Optional dynamic control for optimization of dose distribution. 4. Clinical register: Storage and ready daily access of treatment status for each patient, including diagnostic and treatment data, patient administrative information and data from laboratory investigations and other treatment modalities.
Suntharalingam et al.35 discuss the potential for use of computers and their mass storage media in radiation therapy planning and delivery. They can provide the means for integration of the large amount of data that is related to treatment planning, delivery, and follow-up. Dickof et al.7 describes an integrated radiotherapy computing system that provides for transfer of patient contour data from various sources to the treatment planning system, and its output to the verify and record system. This computing system interfaces an active tumor registry system. Fox et al.11 describes acomputer based information system that schedules patient appointments at each patient care activity in the radiation oncology department. Video displays in each service area provide information on patients scheduled for that service. The system assists personnel in reacting to improve patient flow. Kijewski23 describes an information management system for a radiation therapy facility in which the user responds to menus and prompts in developing command procedures for their application. This shifting of some of the programming burden to the user on an interactive basis permitted reduction of the system programming effort. The information system application areas include: administration, patient registration, biological laboratory, physics, clinical studies, program documentation, machine calibration, radiation safety, patient dosimetry, and reprint library. It has been used on a daily basis for patient records, dosimetry, treatment planning, patient studies, and bibliographical retrieval. The ICRU has issued a report19 on the use of computers in treatment planning and recording and documentation procedures in external beam x-ray and electron radiotherapy. It includes recommendations on quality assurance of the computer system. Figure 10-13 is from this report and illustrates data flow within a department of radiotherapy. Figure 10-14 shows the pathways for beam data acquisition for treatment planning and Figure 10-15 shows patient data acquisition for treatment planning.
REFERENCES
185
FIGURE 10-11 . Information flow in the radiotherapy clinic as it is presented by the Nordic CART project. (from Helax, Box 1704, S-75147 U p psala, Sweden).
REFERENCES 8. 1. ACR-NEMA: Digital imaging and communications standard: July 1, 1985. 2. ACR-NEMA: Digital imaging and communications standards committee; Application guide: Minimum requirements for compatibility with the ACR-NEMA Digital Imaging and Communications Standard, 1987. 3. Cheong VE, RA Hirschheim: Local area nerworks. New York, Wiley, 1983, p 39. 4. Coats JC: Current developments in radiation treatment of cancer by integration of computer technology and the Clinac linear accelerator: Varian Associates, Palo Alto, CA 94303, February 1985. 5. Dahlin H: Program CART-A Nordic challenge in medical computing: Eighth International Conference on the Use of Computers in Radiotherapy. IEEE 84 CH 2048-7. Toronto, July 9-12,1984, pp 541-543. 6. Dickof P, P Morris, D Getz: Vrx: A verify-record system for radiotherapy. Med Phys 11: 525-527, 1984. 7. Dickof P, P Morris: An integrated radiotherapy computing environment: Eighth International Conference on the Use of Comput-
9. 10.
11.
12.
13. 14.
ers in Radiotherapy. IEEE 84 CH 2048-7. Toronto, July 9-12, 1984 pp 549-552. Dorn WL, HP Heilmann, B Bosau: Automatic verification and reporting in percutaneous megavoltage therapy. I: Experience with the Vericord system. Medical Mundi 26(3):150-155, 1981. Enviro-med: A study of the automation of radiation therapy treatment machines. 1972. Ethernet, a local area network, data link layer and physical layer specifications, version 2.0: Stamford CT,Xerox Corp., November 1982. Fox S, JM Hanson, BD Stoskoph: An integrated scheduling and patient flow management system for radiation oncology: Eighth International Conference on the Use of Computers in Radiotherapy. IEEE 84 CH 2048-7. Toronto, July 9-12,1984, pp 553-557. Fredrickson DH, CJ Karzmark, DC Rust, M Tuschman: Experience with computer monitoring, verification and record keeping in radiotherapy procedures using a Clinac 4. Int J Radiat Oncol Biol Phys 5 4 1 5 4 1 8 , 1979. Gibson R: Energy interlock system for a linear accelerator. U.S. Patent 4,347,547, 1982. Grant W 111, J Ames, PR Almond: Evaluation of the Therac 6 linear accelerator for radiation therapy. Med Phys 5:448-450, 1978.
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SERVICE STATION CENTRAL Magtape
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2 16 MB
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. Video
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FIGURE 10-12 Example of treatment management system configuration with Ethernet, providing3D mapping of patient anatomy and 3D dose computation, including beam's eye view (from Helax).
15. Greene D: Linear accelerators for radiation therapy. Adam Hilger, Boston, 1986; p 194. 16. Horii SC: The ACR-NEMA Standards: A tutorial on their structure and use, in RL Arenson, RM Friedenberg, (Ed): Computer applications to assist radiology. Symposia Foundation, 1990, pp 405422. 17. IEC 62C: Proposal for new work: Electronic data exchange format for radiotherapy equipment: Brace JA, London NW3 2QG, Royal Free Hospital: Letter, February 22, 1988. 18. IEC: Draft Publication 601-2-1 revision, Medical electrical equipment, Part two; Particular requirements for the safety of medical electron accelerators in the range 1 MeV to 50 MeV, Section five: Radiation safety requirements, p 48, 1990. 19. ICRU Report 42: Use of computers in external beam radiotherapy procedures with high-energy photons and electrons. International Commission on Radiation Units and Measurements, 7910 Woodmont Ave., Bethesda, MD 20812, pp 1-70,1987.
20. Joint U.S./Scandinavian symposium on future directions of computer aided radiotherapy. (Announcement): Med Phys 15478, 1988. 21. Kartha PK, A Chung-bin, T Wachtor, F Hendrickson: Accuracy in radiotherapy treatment. Int J Rad Oncol Biol Phys 2:797-799, 1977. 22. Karzmark, CJ: Procedural and operator error aspects of radiation accidents in radiotherapy. Int J Rad Oncol Biol Phys 13: 15991601,1987. 23. Kijewski PK: The role of the end user in developing computerized information management applications: Eighth Int. Conf. Use of Computers in Radiotherapy. IEEE 84 CH 2048-7. Toronto, July 9-12,1984,524-528. 24. Kraus HK, A Hess, R Schmidt, KH Hubener: Verification and recording of percutaneous radiation therapy with the Philips SL 75/20 linear accelerator using Vericard S2 system. Medicamundi 34: 34-39,1989.
187
REFERENCES
Identity History Status Diagnosis Follow-Up
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. Beam data acquisitionand input to a treatment planning system (from Ref. 19).
188
CHAPTER 10. SAFETY INTERLOCKING
MECHANICAL DEVICES
I PATIENT DATA ACQUISITION ] FIGURE 10-15
matrix
. Patient data acquisition and input to a treatment planning system (from Ref. 19).
25. Loyd M, H Chow, J Laxton, I Rosen, R Lane: Dose delivery error detection by a computer-controlled linear accelertor. Med Phys 16: 137-139,1989. 26. Miller RW, J van de Geijn: Modification of the fault logic circuit of a high-energy linear accelerator to accommodate selectively coded large-field wedges. Med Phys 14: 262-264,1987. 27. Mohan R, R Caley: Standardizationof therapy machine interface for treatment monitoring. Int J Rad Oncol Biol Phys 9: 12251229,1983. 28. Mohan R, KD Poidmaniczky R Caley, A Lapidus, JS Laughlin: A computerizedrecordand verify system forradiation treatments. Inter J Rad Oncol Biol Phys 10: 1975-1985, 1984. 29. O'Brien PF, RB Barnett, HB Michaels, RA Siwek: Measurements in high-intensitybeams for medical linear accelerators.Med Phys 14: 1067-1070, 1987. 30. Ragan DP: Computer management systems in clinical facilities. Short Course MPSC10. Proceedings of the World Congress on Medical Physics and Biological Engineering. Phys Med Biol33 Suppl. 1 5 , 1988. 31. Rosen 11, MD Stevens, JW Somers, RG Lane, CA Kelsey: Computer interface for a linear accelerator. Med Phys 7:68-69, 1980. 32. Rosenbloom ME, LJ Killick, Bentley RE: Verification and recording of radiotherapy treatments using a small computer. Br J Radio1 50: 637-644, 1977.
33. Stemick ES, JR Berry, B Curran, SA Loomis: Real-time computer verification for radiation therapy treatment machines. Radiology 131:258-262, 1979. 34. Stevens MD, I1 Rosen, RG Lane: Satellite digital display for the Clinac 18. Med Phys 4:454-455,1977. 35. Suntharalingam N, M Goitein, PK Kijewski, J Purdy, G Svensson: Treatment planning and delivery. Cancer Treatment Symp 1:27-33, 1984. 36. Swanson WD: Radiological safety aspects of the operation of electron linear accelerators. IAEA Report 188 1-327, 1979. 37. Taaffe JL, M Kaldis, J Gahm, et al.: Q-RSTAR workstation and system: Technical overview, 317-323 and Q-RSTAR digital image management and transmission, 432-438 in RL Arenson, RM Friedenberg,(Ed): Computer applicationsto assist radiology. Symposia Foundation, pp 432-438, 1990. 38. Tofts PS, RW Cranage (Eds): Networks and image handling. 2 Low Ousegate, York YO1 lQU, UK, The Institute of Physical Sciences in Medicine, 1987. 39. Weinhous MS, JA Purdy, CO Granda: Testing of a medical linear accelerator's computer-control system. Med Phys 17:95-102, 1990.
C H A P T E R
Multi-X-Ray Energy Accelerators
DESIGN CHALLENGES CLINICAL NEED Radiation oncologists prefer use of a low energy x-ray mode for the majority of patients, but a widely separated high energy x-ray mode for about one-fourth of patients, and an electron mode for about one-eigth of patients. A low energy x-ray beam of 4 or 6 MV provides a desirable depth dose distribution for treatment of tumors at moderate depth (e.g., in the head and neck, breast, and lymphatics). The 6-MV depth dose distribution reaches a maximum at 1.5 cm below the surface, about 67% at 10 cm depth, and with parallel opposed fields the dose distribution can be quite uniform through thinner sections of the body. A high energy x-ray mode provides a clear advantage in treating tumors in thick sections of the body. Examples of appropriate clinical targets for such high energy mode are tumors of the prostate, urinary bladder, cervix, esophagus, lung, and sites deep in the brain. High energy x-ray mode is also useful in protecting bone near the skin such as in treating the nasopharanx and associated lymph system while protecting the mandible. The 18-MV depth dose reaches maximum at about 3.5 cm below the surface, about 78% at 10 cm depth and with parallel opposed fields the dose distribution can be quite uniform throughout thicker sections of the body. By combinations of treatments with such low and high x-ray energy beams, as well as electron boost fields, dose distributions can be more precisely tailored to the individual patient. With this range of modalities in a single accelerator, moving the patient from machine to machine can be avoided. This can expedite patient throughput by reducing total set-up time and avoid proliferation of positional imprecision from the additional patient setups that would otherwise be required. It has been a challenge to accelerator designers to provide all these capabilities in the same machine and still achieve uncompromised performance and beam stability for optimal patient outcomes. Much of the discussion in this chapter applies as well to single x-ray energy machines but multi-x-ray energy design is more demanding.
PERFORMANCE REQUIREMENTS One demanding requirement for the machine designer is to achieve a high dose rate flattened to the corners of a quite large field at low energy such as 6 MV and also to supply a widely separated high energy x-ray mode (e.g., 18 MV) with stable dose distribution in both modes over all gantry angles. For patient comfort and reduced probability of patient motion, as well as to allow for absorption in wedge filters and for treatment at extended distances, high dose rate such as 400 or more cGy lmin at 100 cm SAD should be available at both energies. Ideally, the distribution of dose over all field sizes from small to large should be flat at all clinical depths and should remain flat with rotation of the gantry and beam limiting device. This requires that the electron beam be formed, accelerated, aimed, and controlled in ways that will minimize potential contributions to instability of the final treatment beam. Flat distribution at all depths requires enough excess beam energy at the x-ray target so that an optimal flattening filter construction (for shaping the x-ray spectrum over the field) can be used and still achieve desired depth dose. The ICRU Report No. 244 states that at the best level of current practice, the uncertainty in calibration of dose with a secondary standard ionization chamber in a phantom is k2.5 at 95 percent confidence limit. This is exclusive of uncertainties related to the treatment machine. Even assuming complete randomness in summation of errors, this leaves very little room for machine performance tolerances if an overall dose precision of 2 5 percent is to be met. Both the constancy of machine dose versus calibration standard and constancy of dose spatial distribution versus time are essential. Define gamma as the slope of the curve of tumor control probability (TCP) versus dose at 50 percent TCP. Assume a patient with a TCP of 50 percent and gamma of 3. An underdose of 5 percent over the full-treatment course will reduce the TCP by about 15 percentage points to 35 percent for this patient. Similarly, an overdose of 5 percent may increase the probability of severe damage to normal tissue from 5 to perhaps 9 percent. Also, if the excess dose region of a 3 percent asymmetric field is in the region of a critical organ for
11
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CHAPTER 11. MULTI-X-RAY ENERGY ACCLERATORS
the course of therapy, the excess dose can increase the probability of severe injury to that organ. Perhaps even more important than beam symmetry is the requirement for beam stability (e. g., avoidance of variations in symmetrical flatness and penetrative quality) and this need affects the selection among alternative technical approaches in designing a medical accelerator. The design should anticipate changes in performance of individual items over the life of the machine and provide an adequate margin of performance. In addition to providing a stable beam over all equipment orientations, the machine must be reliable, safe, convenient to use, and readily serviced. The machine must be compact for convenient patient setup and this puts severe spatial constraints on the permissible size and arrangement of major components.
ELECTRON BEAM DURING ACCELERATION In order to reduce variation of beam symmetry with time and equipment position, the variation of the convergence angles of the electron rays at the x-ray target should be minimized. Similarly, the variation in position of the centroid of the electron beam at the x-ray target should be minimized. The amount of chaos of the electron rays making up the beam is called transverse emittance. It is the product of displacement and divergence of the electron rays making up the electron beam. Minimizing the transverse emittance makes it easier to maintain the distribution of electron ray angles and positions in the beam spot at the x-ray target. This need for maintaining a precise beam spot distribution affects the choice of electron gun design and the method of energy switching, as well as methods for steering and beam transport. After initial bunching, the electrons should pass through the middle of each cavity near or at the peak of the oscillating rf electric field to optimize use of rf power and minimize energy spread.
ENERGY STABILITY The IEC3 Publication 977 suggests a tolerance value for stability of depth dose in electron mode which corresponds to t1 percent energy stability at energies above 10 MeV. That is, the deviation of the mean of the energy spectrum transmitted by the energy slit should be limited to 2 1 percent. This is facilitated by accelerating a broad fairly uniform energy spectrum (e.g., 20 percent) in electron mode and using a narrow energy slit such as 6 percent to select out only a portion of this spectrum. Some machines employ a 6 percent energy slit (i.e., 2 3 percent), which is narrow enough to limit variation of the mean energy to less than + 1 percent. Some machines reported in the literature11 had energy slits as wide as 16 percent in older designs and have slits much wider than 6 percent in more current designs. One way to meet this energy stability criterion is to use a quite narrow energy slit in the bend magnet. However, this can
limit the dose rate in the x-ray mode if the same slit is used and if the accelerated beam energy spread is not sufficiently narrow. If high dose rate is to be achieved at low x-ray energy fully flattened to the comers of a large x-ray field by a sufficiently thick flattening filter, the percentage beam current transmission through the bend magnet system to the x-ray target must be relatively high. To pass through a narrow energy slit this requires that the gun inject a beam with low transverse emittance into the accelerator guide and that electrons ride the accelerating wave in such a way as to avoid instabilities and increases in energy spread. Also, high beam transmission through the electron beam collimator and energy slit permits thinner, lighter weight shielding in the radiation head because less beam impinges on these components to produce stray radiation. On the other hand, if a separate wider energy slit is used for x-ray mode, it will permit larger variations in mean energy at the x-ray target if the beam from the accelerator structure is not stable in energy. This can result in larger variations in field flatness (shape of dose profile across the field), unless other means are used to stabilize this shape. It is important to place the energy slit at the point in the bend magnet where initially parallel electron rays cross over, producing a radial focus for good energy discrimination (see Chap. 7). In order to avoid adding shielding at the top of the radiation head, with a consequent increase in isocenter height, the energy slit is placed elsewhere in some machines. Mixing of spatial and energy variations can then result, complicating feedback control of field symmetry and beam energy. Variations in the mean energy of the beam transmitted by the energy slit in a doubly achromatic bend magnet will not affect x-ray field symmetry but will affect the shape of the dose profile across the field, increasing or decreasing the dose rate on the beam axis relative to the periphery of the field. As Sutherlands pointed out, this can also cause a dosimeter calibration error in systems where the ion chamber dose rate electrodes respond to the whole field and, hence, primarily to its periphery, whereas external detectors for calibration are used on beam axis. Feedback from dose rate can be used to control energy to avoid this error. Variations in mean energy also make small variations in relative dose at depth, about V4 percent at 10 cm depth per 1 percent change in mean energy around 6 MeV. If the beam energy spectrum shape is quite peaked or curved, changes in the position of this distribution in the energy slit will cause variations in the mean energy of the energy band transmitted by the slit. If the energy spectrumis quite broad and relatively flat over the energy slit range, variations in the mean energy out of the accelerator guide will be reduced in the beam transmitted by the energy slit.
DOSE SPATIAL DISTRIBUTION AND CALIBRATION IN INITIAL SECONDS The beam energy, dose distribution, and dose calibration should be within tight tolerances even during the initial seconds
191
EQUIPMENT DESIGN ALTERNATIVES
of each portal treatment. In the future, conformal therapy may call for say 10-port treatments of 5 s each. A 10 percent error in field flatness through just the first second could add two percentage points of error to the steady-state value of flatness. Assuming a tumor dose of 200 cGy, a 0.1-cGy round-off error or a depth dose error in the 20 cGy delivered by each of 10 portals could add another 0.5 percent dose error. These errors could become significant if they add systematically. A similar rationale applies with hyperfractionation where there are more beam initiations in a course of treatment. Movement of the electron beam centroid at the x-ray target in the initial second(s) can contribute to poor spatial resolution (geometric unsharpness) in portal images, especially pretreatment images taken with geometric enlargement. Such pretreatment open fields are taken in a time of order 1 s or less to minimize patient dose outside the treatment field. Fast rise and precisely controllable dose rate during treatment can be essential in dynamic radiotherapy. Examples are dynamic wedge with moving jaws, dynamic conformal arc with moving leaves of multileaf collimator, dynamic compensation, and scanned slit. These can be facilitated by avoiding use of components with long thermal time constants (e.g., long anode magnetron).
EQUIPMENT DESIGN ALTERNATIVES In the following discussion, alternative designs for dual x-ray energy accelerators will be compared, explaining the reasons for particular design choices. Emphasis will be placed on minimization of treatment beam instabilities and on having adequate performance reserve to ensure long machine life with minimal maintenance. Although dual x-ray energy linacs had been available from about 1970 [Varian Clinac 35; Philips (MEL) SL751201 (see Appendix B), they had limitations that inhibited widespread acceptance by the radiotherapy community. Since 1983, manufacturers have been providing dual x-ray energy linacs, which more properly fulfill the needs of radiotherapy. These machines are economically justifiable and provide both low and high energy x-ray modes as well as a full range of electron mode energies. The machines are compact, permitting installation in conventional size treatment rooms.
MICROWAVE POWER SOURCE-KLYSTRON VERSUS MAGNETRON A klystron is used in most dual high energy x-ray energy machines because it provides high pulse power conservatively and because it operates as an amplifier, being driven by a separate frequency-stabilized oscillator. Thus, the phase and amplitude of the rf output power are highly independent of reflections from the accelerator guide. A klystron is preferred
for higher energy multimode accelerators because the high pulse power permits a shorter accelerator guide for better beam stability. An alternative would be to use a magnetron. This is conventional in low energy machines but is more difficult at 18 MV or higher x-ray energy. The magnetron is an oscillator and its output power, phase, and frequency are highly dependent on reflection from the accelerator guide. Also, because of its compactness there is a limit to the maximum pulse and average power that can be obtained reliably from amagnetron. Attempts to exceed this limit can result in increased arcing, increased frequency moding, and decreased operating life. Because the pulse power is limited, achieving high energies with a magnetron requires a longer accelerator guide, with corresponding potential for increased beam instabilities. Because magnetron average power is limited, it is more difficult to achieve high dose rate over large fields fully flattened to the comers in low energy x-ray mode (such as 6 MV) in a dual x-ray energy machine.
ELECTRON GUN-TRIODE
VERSUS DIODE
A triode gun employs three electrodes+athode, grid, and anode. Triode guns typically employ an impregnated cathode. Operation in space charge limited mode provides a virtual cathode, which minimizes transverse emittance of the electron beam from the gun. In electron mode, the gridlcathode voltage is automatically limited to a low or negative value to ensure that abnormally high, hence unsafe, electron current cannot be emitted. Dose per pulse can be regulated by the control grid beam pulse length, clipping the end (not the beginning) of the beam pulse so as to maintain constant beam acceptance by the accelerating field. The ability to vary dose rate rapidly and precisely is especially important in arc therapy with constant gantry rotation rate. Care must be taken in designing the grid, whether it be of the intercepting or nonintercepting type, to ensure that it does not increase the beam emittance excessively, such as by local transverse electric fields inside the beam. Instead of a triode, a diode gun may be used, employing only a cathode and an anode. A thoriated tungsten wire spiral or circular type directly heated cathode may be used. Such cathodes operate at much higher temperatures than impregnated cathodes. Because of local electric and magnetic fields at the wire, as well as their thermal energy, the electrons leave over a spread in angles, increasing the transverse emittance. Also, a wire spiral cathode can distort with time and temperature.The emission current can be varied in temperature limited mode by varying the heater current. However, this can destroy the space charge virtual cathode and increase beam emittance still further. Instead of using temperature control, the diode gun emission current may be varied by variation of the cathodelanode voltage. This can lead to instabilitiesand enlarged energy spread
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CHAPTER 11. MULTI-X-RAY ENERGY ACCLERATORS
because of variation of injection energy and, hence, phase of the electron bunch on the wave in the accelerator guide.
ACCELERATOR GUIDE-TRAVELING VERSUS STANDING WAVE
WAVE
One type of accelerating structure is traveling-wave (TW) guide. The rf power is coupled from one cavity to the next through a hole on the axis. The hole diameter must be quite large, about one-fifth the cavity diameter, in order to provide enough rf power flow. Hence, the rf electric field lines curve away from axis to the hole periphery, providing a reduced longitudinal component for acceleration of the electron bunch on the axis. Optimum acceleration is obtained with three cavities per wavelength in TW guide instead of the two cavities per wavelength of standing-wave (SW) guide. The increased number of cavity walls absorbs more rf power. The net result is that for a given maximum no-load beam energy and rf pulse power from the microwave source, a TW guide must be about 1.7 times as long as a SW guide. With a drum type gantry, the accelerator guide can extend back through the drum; hence, this extra length is permissible. However, in the low energy x-ray mode, the electrons are at low energy over a much longer distance and, hence, are more sensitive to destabilizing transverse forces. For exam-
Acceleration Portion
Energy Switch
ple, unless carefully shielded, stray magnetic fields at the beam can vary with angular position of the gantry. Because of the much weaker coupling between cavities in a TW guide, temperature variations can cause much greater changes in phase over the guide length between the rf field and the electron bunch, thereby contributing to instability of beam energy. This can be especially important during the initial seconds of beam on time. In an SW guide, the peaks of the oscillating sine wave electric field are located in each accelerating cavity and the nodes of this wave are in the coupling cavities. Since the electric field in these coupling cavities is near zero, they can be made very small and can be placed on the side of the guide or between accelerating cavities with little attendant power loss. In an SW type accelerator guide employing a microwave energy switch, the lengths of the first few cavities are slightly foreshortened. Electrons injected from the gun pass through the midplane of each cavity at a time when the oscillating rf electric field is rising (see Fig. 11-la, and b). Early electrons see lower amplitude and late electrons see higher amplitude electric field as it rises sinusoidally in time, hence coalescing to a tight bunch as they gain velocity rapidly and drift back toward the crest of the electric field wave. In the remainder of the guide, the electron bunch passes through
Acceleration Portion
0" From
Electron Bunch
Phase Spread of Bunch
FIGURE 11-1 . Electron bunch phase with energy switch. The dashed wave is 180" of phase later in time than the solid wave. (a) 18 MV x-ray mode and (h) 6 MV x-ray mode.
EQUIPMENT DESIGIN ALTERNATIVES
the midplane of each cavity when the oscillating rf electric field has reached approximately its peak. For example, with cavities at 5 cm intervals, the electron bunches are 10 cm apart. This SW guide design with microwave energy switch has several beneficial attributes. 1. The beam energy spread is minimized. This permits use of a narrow energy slit in the bend magnet to define the centroid of the transmitted energy spectrum with better precision and stability and reduces stray radiation due to beam loss in the bend magnet chamber. 2. The sensitivity of the electron bunch energy to variations in phase of the rf electric field is minimized. 3. The requirement for rf pulse power to establish the accelerating electric field is minimized, easing the demand on the microwave source (i.e., copper losses are minimized, saving more rf power to accelerate the beam). Instead of off-axis (side-coupled) design, an alternate way to build an SW accelerator guide is to make the coupling cavities very short and put them on axis, alternating with each accelerating cavity (bi-modal) or each pair of accelerating cavities (tri-modal).' This permits using a smaller diameter solenoid if one is used to confine the beam diameter. The coupling cavities provide no acceleration, yet take up axial space. Hence, for the same overall accelerator guide length and basic cavity shape and the same maximum electron energy, somewhat greater rf pulse power is required from the microwave power source with on-axis coupling cavities. Also, the beam can excite repelling fields in these cavities, disturbing the phase distribution at high beam currents. An SW type accelerator structure (side coupled, bi-modal, tri-modal) has the following important characteristics, which are superior to a TW guide. 1. The shape of the accelerating cavity (and especially near the axis) can be optimized to concentrate the rf electric field in time and space at the electron bunch to achieve a given energy gain per cavity with minimum rf pulse power loss to the cavity surface-and independently the coupling cavity can be designed for optimum transfer of excitation rf power between accelerating cavities. The shunt impedance is thereby maximized. That is, efficiency of conversion of rfpower to accelerating electric field is maximized, permitting shorter accelerator guide length. 2. The accelerating cavities can be tightly coupled so that they maintain a close rf phase relationship to each other and the traveling electron bunch. 3. The use of half-wave accelerating cavities results in the guide being inherently mode stable, suppressing potential oscillation in nearest neighbor resonant modes, which are about 1 percent away in frequency. Feeding the rf power to the accelerating guide near its midlength instead of at one end contributes to this mode stability. Avoidance of
193
excitation of neighboring resonant modes contributes to stability of beam characteristics.
SWITCHING FROM HIGH TO LOW ENERGY IN A TRAVELING WAVE GUIDE Special techniques such as an energy switch are not needed in a TW guide. The rf electric field can be high in the first portion and can remain high in the second portion of a TW guide for a high energy x-ray mode or can taper down to a low (or reverse) field over the length of the second portion for a low energy x-ray mode, with the electrons well bunched and riding the crest (or valley) of the wave in either case. The switching from a high to a low x-ray energy mode can be accomplished while keeping the microwave source power and buncher electric field constant simply by increasing the beam current, so it progressively extracts an increasing fraction of the rfpower flow down the TW guide. This is called beam loading. The beam energy can also be varied by detuning the rf source, so that the electron bunch drifts in phase over the accelerating rf sine wave field. The rfpower travels only in the forward direction through a TW guide. Radio frequency feedback through an external circuit permits opening up the axial holes for tighter coupling, in order to reduce phase instability. However, this reduces efficiency of conversion of rf power to an accelerating electric field and the system still ends up about 4 to 10 times less stable in phase than an SW guide, with corresponding increased energy instability, especially with a magnetron as microwave power source.
SWITCHING FROM HIGH TO LOW X-RAY ENERGY IN A STANDING WAVE GUIDE Obtaining especially high dose rate at 4 or 6 MV in a multi-xray energy accelerator requires optimal design. Such high dose rate can be important for some applications, such as physiological gating of the treatment beam to compensate for organ motion and is facilitated by the SW guide design. There are various techniques used for varying the output electron energy of an SW linac. The simplest method is either to vary the rf power into the accelerator or vary the injection beam current and, hence, beam loading. However, this approach will provide only a narrow range of energy variation beyond which the energy spectrum will rapidly degrade (see Fig. 11-2). Another technique is to detune the rf source frequency and/or detune part of the accelerating cavities. In this case, however, the energy range will be limited by the lack of stability of the system as it is detuned from the stable condition. One practical approach to vary the energy over a wide range is to cascade sections of the accelerator that are independently excited from a common rf source with independent control of the phase and amplitude. However, the rf system and
194
CHAPTER 11. MULTI-X-RAY ENERGY ACCLERATORS
Energy (MeV) Beam Loading: a, heavy; b, optimal; c, light
FIGURE 11-2 . Deterioration of energy spectrum of standing wave accelerator optimized for 10 MeV when energy is varied siniply by varying rf power into the accelerator structure or varying the injected beam current and, hence beam loading.
microwave structure of such an accelerator can get bulky, complicated, and expensive. Another method has been tried to achieve the equivalent of a multiple section accelerator, namely, the double-pass single-section linac. It uses a mechanically movable 180" reflecting magnet to obtain the correct phase variation between the beam passes. Energy can then be controlled by adjusting the phase of the bunch on its return relative to the accelerating field. This technique is complicated and moreover, the degradation of the spectrum and the lack of system stability make it difficult to implement the idea. It is especially difficult to obtain high dose rate over large fields fully flattened to the comers at low energy such as 6 MV because the poor energy spectrum causes large loss of beam current on the energy slits in the bend magnets and also large leakage radiation. For the following discussion, the guide length is arbitrarily divided into a first portion in which the major function is to bunch the electrons and position them on the accelerating wave, and a second portion in which the primary function is to accelerate the electrons while minimizing their output energy spread. In a high energy x-ray mode, the second portion of the guide must be at a high rf field, with the electrons tightly bunched and riding near or on the crest of the rf electric field in order to minimize energy spread, energy instabilities, and demands for pulse power from the microwave source. In a low energy x-ray mode, the electrons must experience a low net average rf electric field in the second portion of guide but still be properly bunched and accelerated in the first portion. There are two fundamentally different ways to do this in an SW guide. 1.
One method is to use a compact microwave switch to reduce the rf power to the second portion of the guide, thereby reducing the rf electric field so the electron bunch can ride the crest of a reduced wave. This method is efficient in use of microwave pulse power and in attaining minimum beam energy spread. (In Figure 11 -la and b, the
FIGURE 11-3 . Microwave energy switch types for an SW guide. (a) The noncontact method, (b)the single side cavity shorting method, and (c) the double side cavity shortinglopening method.
195
EQUIPMENT DESIGN ALTERNATIVES
dashed wave represents the rf field during the time that the cavities have oscillated into reverse phase, 180"later in the time than the solid wave.) The microwave energy switch may employ a noncontact method9 (Fig. 11-3a) or a shorting method5'10 (Fig. 11-3b and c). The noncontact method is preferred because it permits transmission of enough rf power to maintain a desired low accelerating field in the second portion of the guide. The shorting method in a single coupling cavity results in the beam exciting the second portion of the guide, building up a decelerating rf electric field over a second filling time. This reduces the pulse time in which acceptable beam energy spread is achieved, resulting in reduced x-ray dose rate. Two side coupling cavities diametrically opposite each other can be used to couple a single set of on-axis cavities, using higher coupling coefficient slots for one side cavity, lower coupling coefficient slots for the other side cavity.5 By shorting one side cavity and unshorting the other, and vice versa, the accelerating field in the second portion of the accelerator guide can be switched from high to low and vice versa from low to high (see Fig. 1I - 3 ~ ) .
There were reliability problems with early versions of microwave energy switches, related to the use of apart that was moved in vacuum. Improvements have been achieved via redesign. A second method is to design the first portion of the guide (the buncher) so that the electron bunch is formed even at a reduced level of peak rf electric field and then simply change rf power to change beam energy (see Fig. 11-4). This has been called a broadband buncher. It avoids the use of a part that moves in vacuum. In order to form an electron bunch around a usable phase angle at widely different rf peak electric field levels, the electrons would spend considerable time slowly gaining energy at a synchronous phase angle well forward of the crest in significantly foreshortened cavities, gain only limited energy while they bunch, and then be shifted in phase to near the crest of the rf electric field for the rest of the guide. This technique is wasteful of rf power in the buncher region, so it requires higher peak power from the microwave source to obtain a given maximum x-ray energy in a given guide length. Figure 11-4 shows uniform field over the guide length. Somewhat better results could be obtained by using
/
Cell Number Phase Transition Cavity Broadband Bunching Portion
- 18" From Crest
I
-
Acceleration Portion Electron Bunch Phase
(+I h
(-)
of Bunch
FIGURE 11-4 . Electron bunch phase with broadband buncher. The dashed wave is 180" of phase later in time than the solid wave. (a) 18 MV x-ray mode and (b) 6 MV x-ray mode.
196
CHAPTER 11. MULTI-X-RAY ENERGY ACCLERATORS
Phase of V, Ahead of V, By 0
\1-b +.
Net Field V, Shifts Ahead of Electron Bunch By $
(
+e!+ I
I
I
-
! Unloaded Field Generated By Klystron or Magnetron
Bunch Centroid Net Accelerating Field Beam Energy Spread
I
Opposing Field Generated By High Electron Beam Current
FIGURE 11-5
I I
-
. Effect of beam bunch phase error on phase of net accelerating field and on energy spread with heavy beam loading.
a lower field in the first few cavities than in the rest of the guide, but this is even more wasteful of rfpower. Changing x-ray energy over a 3-1 range, from 18 to 6 MV, will require a compromise in phase angle of the electron bunch, being too far forward on the rf electric field wave at 18 MV and too far behind at 6 MV. This will result in increased beam energy spread, hence reduced transrnission through the energy slit to the x-ray target in 6-MV mode, hence reduced dose rate for equal maximum field size flattened fully to the comers. Running the electron bunch at a phase angle off the crest results in increased sensitivity of electron energy to variations in electron beam current loading and rf input power.' Figure 11-5 illustrates this point. For example, a sudden increase in beam current increases VB,decreasing the net accelerating field and shifting the phase of its crest from 0 to with respect to the electron bunch. This perturbation (shift of net accelerating field relative to the electron bunch) damps out with a time constant of about 0.01 FS in an SW guide. In a TW guide, even
+
with rf feedback, this effect is much larger because the damping time constant is about 0.5 FS. In order to achieve useful (but still marginal) dose rate in low energy x-ray mode with a broadband buncher in an SW guide, it may be necessary to make compromises such as: 1. Use of a thinner flattening filter, hence not flattening to the comers of large fields (see Fig. 11-6). 2. Use of larger electron beam collimator holes. This permits transmission of greater transverse emittance and, hence, produces larger beam divergence angles at the x-ray target. This can exacerbate beam symmetry instabilities. 3. Use of wider energy slit in the bend magnet system. This can permit larger instabilities of depth dose curve in electron mode and larger instabilities of symmetrical field flatness in x-ray mode. 4. Injection and acceleration of excess beam current and wasting much of it on collimator holes and energy slit. This exacerbates energy instability due to off-phase interaction with the rfaccelerating electric field. Also, the requirement
EQUIPMENT DESIGN ALTERNATIVES
/
/
40 cm x 40 cm Maximum Square Field
-----
FIGURE 11-6 . Flattened circle and size of square field flattened to corners for various thickness flattening filters.
for increased radiation shielding can affect the size and weight of the radiation head and, consequently, the height of the isocenter for clearance from the floor or, alternatively, the location of the energy slit. There are other methods that could be used to provide operation at two widely separated x-ray energies with an SW guide without using an energy switch. For example, with appropriately tuned guide cavities, the guide could be powered at one frequency to produce a uniform rf electric field over the guide length for a high energy x-ray mode, and at a different but closely neighboring frequency to produce a cosine-like distribution of rf electric field for low energy x-ray mode, maintaining high field in the buncher in both cases. However, such techniques are potentially unstable, being very sensitive to thermal and frequency variations. The fundamental problem still remains. The electron bunch is displaced somewhat from the rf electric field crest in the low energy x-ray mode.
BEAM L O A D I N G For an electron bunch of very short length, which rides the crest of the fundamental space harmonic of the axial field in the accelerator structure, load lines of beam energy Vversus beam
197
current i can be calculated using the equations of Table 11-1, which are derived from Lapostolle6 and Neal.' Examples of load lines are shown in Chap. 4. In essence, the net energy V = Vo-AV, where Vo is the energy gain of an electron on the crest at zero beam current and AVis the reduction in energy caused by the beam load current i. Load lines are only a small part of the story in estimating maximum dose rate at various energies. Bunching, off-phase induced fields, radial divergence due to space charge at injection, beam focusing, and so forth, are important factors. Computer calculations are required to determine the beam intensity profile at the electron collimator and the energy spectrum of the beam and its intensity profile at the energy slit. Heavy beam loading exacerbates problems related to bunching, off-phase transient induced fields, and resultant spreading of the beam energy spectrum. In T W structures of constant gradient design, the group velocity of the rf wave decreases toward the exit end. That is, the coupling between cavities is much lower toward the exit end than near the entrance in order to tap off successively less power from the preceding cavity. With heavy beamloading the field induced by the beam increases toward the exit end, increasing the tendency for the net accelerating field to slip out of phase from the electron beam bunches, thereby tending to induce beam energy spread and energy instability. Hence, a small change in the temperature of the guide, of the rf source frequency, or of the beam current can cause a tapered distribution of phase error of the rf field with respect to the beam bunch.
NON-CONTACT T Y P E SIDE CAVITY E N E R G Y SWITCH Figure 11-3a shows the simplest structure of a side-coupled SW accelerator consisting of two centerline cavities and one coupling side cavity used as an energy switch. Its equivalent circuit leads to the relationships between coupling factors (KO]and K,,) and accelerating field amplitudes (Eo and E2) in the centerline cavities for the 1~12mode operation.
Since KQ is much larger than unity, the second term is small. Hence, accelerating fields can be inversely varied by varying the coupling factors Kol and K I 2between the accelerating cavities and coupling side cavity. The negative sign indicates that the phase between Eo and E2 is offset by 180". If a longitudinal asymmetry is introduced by lengthening one post and shortening the other (see Fig. 11-3a), a longitudinally asymmetric field will be excited while the resonant frequency is kept constant. With this asymmetric side cavity coupled to the two centerline cavities, the magnetic coupling of the short post will be significantly reduced and the coupling of the longer post side will be increased. By introducing this side cavity at a particular location along a side-coupled accel-
TABLE 11-1
. Equations for load lines
Traveling-wave guide without feedback:
Traveling-wave guide with feedback:
n
Standing-wave guide in - mode: 2
For an SW guide with an energy switch producing unequal fields over length L1 before and L2 after the switch, the effective shunt impedance rlL is substituted for rL, where:
Where (for all guide types) Shunt impedance per unit length Accelerator guide length Resistive attenuation of rf field over length L Coupling coefficient of input coupler Resistive power loss to guide at zero beam Power incident on input coupler Power reflected from input coupler Power transmitted by input coupler Power coupled to beam Power from rf source into "lossless" bridge of TW feedback network Resistive power loss to guide when beam loaded Ratio of first section length to total length Ratio of first section energy to total energy Effective shunt impedance due to (Y and 5
Abbreviation
Units
Mfllm m nepers
199
REFERENCES
TABLE 11-2
. Clinac-1800 accelerator guide design parameters
Accelerator rf length (m) Number of accelerating cavities Frequency (MHz) Effective shunt Impedance (Mfllm)
Qo Coupling factor, kl Output energy (MeV)a Radio frequency power into guide (MW) Load line (MeV)
18 3.8 V = 23.0 - 70i
1.47 28 + V2 2856 102 15,200 0.04 6 1.2 V = 9.8 - 30i
"Shows data for dual x-ray beams with widest energy spread. Other energies are available.
erator, one can vary the accelerating field ratio while maintaining the 7r12 mode resonance. Thus, the accelerating field in the second part of the accelerator guide can be shifted from high to low while maintaining a constant field for bunching in the first part, (see Fig. 11-1). Table 11-2 summarizes the Clinac 1800 accelerator design parameters. The variable coupling side cavity is located between the rf coupler and the output accelerator section. The buncher is designed to be optimized at an accelerating field of 10 MVIm. The injected electrons will be fully bunched and accelerated to 4 MeV at the point where the fif !d is stepped. Figure 1-24 shows the experimental results of the normal)zed output energy spectrum measured both at 6 and 18 MeV fith a magnetic spectrometer using a 1 percent energy slit. The measured fwhm energy spread for both modes is about 3 percent.
SYSTEM FEEDBACK CONTROL PHILOSOPHY Each manufacturer has its own set of fundamental design philosophies. One is to design each subsystem to be stable unto itself, operate at only specifically selected steps of beam energy and dose rate, and then use simple feedback control to stabilize each step. A second approach would be to depend much more heavily on feedback control via computer, using look-up tables for each operating mode. A third approach is to align the equipment well enough so that it can run with minimum feedback control, accelerate excess beam current, and waste an oversized beam diameter and energy spread on shielded beam collimators and energy slit so that deviations in mean beam position and energy are reduced in the beam current reaching the x-ray target.
REFERENCES 1. Arai S, E Katayama, E Tojyo, K Toshida: Beam loading effects in a standing wave accelerator structure. Particle Accelerators 11:103-111, 1980. 2. Bensussan A, DT Tran, D Tronc: Standing wave triperiodic structure for a 10 MeV medical accelerator. Nucl Instr Meth 140:231-235, 1977. 3. IEC Publication 977: Medical electrical equipment. Medical electron accelerators in the range 1 MeV to 50 MeV--Guidelines for functional performance characteristics. IEC, 1 Rue Varembe, Geneva, Switzerland, 1989. 4. InternationalCommissionon Radiationunits and Measurements: Determination of absorbed dose in a patient irradiated by beams of X or gamma rays in radiotherapy procedures. ICRU Sept. 15, 1976,7910Woodmont Ave.; Washington, DC 20014; ReportNo. 24. 5. Kazusa C, M Yoneda: Side coupled standing wave linear accelerator. U. S. Patent 4,746,839, issued 24 May 1988. 6. Lapostolle PM, AL Septier: Linear accelerators, New York, Wiley, 1970. 7. Neal RB (ed.): The Stanford two-mile accelerator, New York, W. A. Benjamin, 1968. 8. Sutherland, WH: Dose monitoring methods in medical linear accelerators. Br J Rndiol43: 864, 1969. 9. Tanabe E, R Hamm: Compact multi-energy electron linear accelerator. Nucl Instr Meth B10/11:871-876, 1985. 10. Uetomi I: Standing wave accelerator. U. S. Patent 4,651,057, issued 1987. 11. Whitham K: 20 MeV S-band standing waveguide. IEEE NS-22 (3), 1328-1333,1975.
C H A P T E R
Patient Support Assembly and Treatment Accessories
PATIENT SUPPORT ASSEMBLY The patient support assembly (PSA), be it treatment table (couch) or treatment chair, fills a central role in supporting, positioning, and facilitating immobilizion of the patient in a readily implemented and comfortable manner. The PSA must be constructed with high standards for safety, accuracy, rigidity, reproducibility, reliability, maintainability, together with flexibility in use for patient treatment and portal imaging. The need for reproducible positioning of the patient during the entire course of radiation therapy, which ensures directing the beam accurately on the target volume, has been demonstrated by Kartha et al.,23 as well as by Marks et al.40 Ideally, one uses the same PSAon the simulator to facilitate duplication of positioning for therapy. Figure 12-1 illustrates an elementary treatment couch for an isocentric linac and emphasizes the geometrical relationship of the linac and treatment couch motions. The couch support must be offset from the isocenter and the couch cantilevered to allow the radiation head to pass underneath the couch in 360" gantry designs (early machines did not rotate 360"). Patient chairs may be attached to the end of the couch for positioning and make use of the coordinate motions of the couch. They are used infrequently, typically for 30 or fewer patients per year, less than 10% of patient load.
PATIENT TABLE SUPPORT TYPES Treatment couches are mounted on a pedestal (see Fig. 12-1) or on a vertical column called a ram (see Fig. 12-2).Couches incorporate motorized control of longitudinal, lateral, vertical, and angularpositions usually via a lightweight,hand-held or ceiling mounted pendant. Such controls may be duplicated on the couch. The longitudinal and lateral motions of the couch top may also be positioned manually. Some couch tops provide an auxilliary rotation about a vertical axis displaced from the isocenter as well as roll and pitch angular motions about lateral
and longitudinal couch axes respectively.21 The couch may incorporate pushbutton controls for room, field, rangefinder, and laser positioning lights, as well as for retractable beam stopper positioning and for emergency machine off. The couch rotates about a vertical axis, passing through the isocenter normal to the gantry axis (see Figs. 12-1). Couch rotation facilitates certain techniques such as stereotactic radiosurgery, wherein multiple arcs of the gantry are spaced over the patient's head by rotating the couch to a number of discrete couch rotation angles16 (see Figs. 2-12 and 2-13). Couch rotation is also used in contemporary breast treatment with complex 3-D positioning techniques.35.57 It is important that this rotation have an easily read angular scale, a secure position lock, and a true vertical axis. The pedestal vertical motion (Fig 12-1) is powered by a contoured scissors-jack mechanism contained in the base that gives uniform lift speed for a constant motor drive speed. The pedestal style couch is largely self-contained above the floor but often has a high minimum height position above the floor, which precludes the use of large SSD values and treatment field sizes. Recently, extended range, double-scissor-jack couches have been introduced with more than lOOcm of travel and having both high and low height capability. This permits large field treatments from both above and below arecumbent supine patient, through large openings in the couch top, eliminating the internal organ shift when changing between the supine and the prone position, a change usually necessitated if vertical couch travel is limited. Usually the ram style couch can be positioned nearer the floor to permit large treatment field sizes and SSD distances as large as 150 cm from above and below the patient. However, such a couch may require a deep pit, up to 6 ft, in the floor with attendant cost, maintenance inconvenience and vulnerability to flooding from both hospital and machine water supply. The ram vertical motion is powered by a long, motor driven lead screw. The couch itself may be detachable from the lifting ram, leaving a clear floor area for special patient setups and to facilitate the use of dosimetry apparatus. Couch motions are
12
202
CHAPTER 12. PATIENT SUPPORT ASSEMBLY AND TREATMENT ACCESSORIES
GANTRY X-RAY TARGET
\
P
CEILING LASER
COLLIMATOR ROTATION
GANTRY ROTATION
DIGITAL POSITION INDICATORS
BEAM CENTRAL AXlS
STAND SlDE LASER
SlDE LASER
I
/ -
AXlS
1 \
/
\\\\A \\\\. \
,
ISOCENTER
TRANSLATIONS
TREATMENT COUCH
.COUCH ROTATION /
HAND CONTROL
FIGURE 12-1 . Schematic view of a treatment unit emphasizing the geometric relationship of the linac and treatment couch motions. A pedestal type couch is illusTrated.
FIGURE 12-2
. Ran1type patient couch. (Courtesy of Philips.)
usually equipped with electrically released brakes. Drive motors are provided with slip clutches to prevent inadvertent application of excessive torques. The isocenter height of a treatment unit is determined by the length of the radiation head on the beam axis and the target-skin distance. Operationally, a low isocenter height is favored for the convenience of technologists, but it may be associated with more restricted, vertical couch travel and collimator clearances. Vertical couch motion is always powered, preferably variable in speed and interlocked with a safety enable bar "deadman" switch on the couch control pendant. If not powered, lateral and longitudinal motions of the couch are provided with locks that facilitate orderly setup procedures. Flexible couch positioning aids both ambulatory and gurney patients in mounting and dismounting or transference to and from the patient couch (table top).
PATIENT SUPPORT ASSEMBLY
Several modifications have been developed to overcome some of the limitations of early couch designs. Gillen et al.13 constructed a vertical couch extender that overcomes the limitations of some couches in permitting the use of large anterior and posterior fields while keeping the patient in the supine position during the entire course of treatment. Biggs and Leong3 describe an attachment for an existing couch to extend the range of SSDs at which patients can be treated with a downward directed beam. It attaches perpendicular to the end of the existing couch and makes use of all the latter's motions. A heavy patient may exert significant stress on the treatment couch, particularly, if supported in extreme positions. The IEC has established positioning range and scale standards, deflection tolerances under specified loading conditions, as well as specifications for the parallelism of vertical motion and rotation axes.21 If the possibility exists that failure of a motorized movement during normal use of the equipment might result in a patient becoming trapped, means shall be provided to permit the release of the patient.21 Simulator couches should incorporate similar standards and tolerances.
TABLE TOP The couch top is often constructed of smooth, hard panel sections, for example, Benelex (Masonite) plastic, or carbon fiber for strength and x-ray transmission which attenuates the treatment beam minimally. An early finding was that deeply upholstered couch tops, although more comfortable for the patient, were unsatisfactory for reliable, reproducible patient positioning. Couch sections are of standard width, typically about 50 cm, of variable lengths, and are removable and interchangeable to best accommodate the treatment beams and patient anatomy. Often, one section of the couch top is constructed of a nylon "tennis racket" grid (covered with Mylar for comfort) or a thin Mylar film to facilitate observation for treatment procedures from below and for port film exposures, as well as for marking the skin from below (through the tennis racket)? Treatment planning must take cognizance of possible patient sag between different CT, simulator, and treatment table top surfaces. When port films are needed, the couch should accommodate film cassette holders for films as large as 14 X 17 in. between rails or the cassette supported from ancillary devices. Some couches incorporate a center spine section, which facilitates the placement of angled posterior fields without attenuation of the beam by the couch rails. Many couch tops may be rotated end-for-end about a central vertical axis. Such a couch is constructed with an outer, dual-side-rail structure on one end and a single center spine on the other end. This feature provides flexibility in the treatment field and portal film orientation. To avoid interference from the couch rail (s), central lesions irradiated with anterior-posterior fields make use of the dual-side-rail end of the couch and lesions employing angled fields often make use of the single center spine end. Accessory rails are usually provided along the
203
couch dual-side-rails, and often along their ends, to attach patient positioning and immobilization devices, as well as field modification accessories and cassette holders. These and other accessories are treated on pages 204, 208. Simulators should be equipped with therapy style couches having similar tops and attachable accessories. However, simulator couches may be fitted with a carbon fiber couch top, a plastic composite featuring strength, lightness, and translucency to diagnostic energy x-rays.19
TREATMENT CHAIR Some lesions may be advantageously treated with the patient seated. These include pituitary and other brain tumors, as well as other head and neck sites and particularly, central lesions of the head. Treatments carried out with the patient seated may be particularly valuable for a small group 'of patients unable to lie down and for treatment areas more accessible with the patient in an upright position. Immobilization of the patient's head using a bite block is often used in conjunction with chair treatments.24 A chair described by Karzmark et a1.,*5.26 accommodates both sitting and standing patients. Drive mechanisms for this chair are located in a floor pit with the chair detachable, leaving a smooth floor for patients wheeled on a gurney. The chair itself is mounted on casters to facilitate its attachment, removal, and storage. The powered motions include horizontal X and Y translations; two concentric rotation axes about the vertical and two SAD positions along the horizontal axis of the treatment beam. A commercial chair is described by Mulkerin and shown in Figure 12-3.4The chair is fixed to the end of the treatment
FIGURE 12-3 . Tkeatment chair shown attached to end of treatment couch. (Courtesy of Varian.)
204
CHAPTER 12. PATIENT SUPPORT ASSEMBLY AND TREATMENT ACCESSORIES
couch with locking pins placed through the dual-rail structural members. Casters on the chair permit technologists to easily roll the chair aside. Such a chair makes use of the couch motions for positioning the patient. The permitted gantry angles for treatment with isocentric linacs will be limited by the chair, particularly, in beamstopper equipped treatment units. Compared with a couch, a treatment chair is more cumbersome, the chair patients will often require more setup time and often, no provision is made for simulation in a chair.
TREATMENT ACCESSORIES Radiotherapy treatment procedures have become more complex and precise in the continuing effort to accurately deliver a prescribed dose to a specified target volume. Increasingly, sophisticated accessories are incorporated into such procedures and the simulation of them. Historically, a small number of treatment centers designed and constructed accessories to fit their own individual needs, often with the assistance of an on-site or a local machine shop. At times, a descriptive publication ensued and other centers adopted or modified the original design. Infrequently, an enterprising manufacturer would produce and market these innovative devices. The adoption of accessories that could enjoy wide usage was slowed by the variability of treatment procedures between centers, by a lack of standardsand agreement as to what was needed, as well as the relatively small market that discouraged commercial production. Over the years, a growing awareness of the importance of accessories in treatment procedures became apparent. This awareness gave rise to the development of new and better accessories.It emphasized the need for their interchangeability between treatment units and simulators, as well as among treatment units themselves. There was also increased agreement as to what accessories were needed and their characteristics. Several manufacturers now cater to these needs independently and increasingly accelerator manufacturers incorporate functional accessories into their patient treatment systems. Accessories may be categorized in terms of their function or their usage. Some accessories, such as shadow blocks, wedge filters, and tissue compensators, function in modifying the radiation beam spatial dose distribution. Others are concerned with patient immobilization, patient alignment, or positioning. Accessories such as wedge filters are mounted in the accessory holder on the radiation head, and others elsewhere on the gantry, the couch, or the floor. Few standards presently exist for accessories. When adopted, they may involve such considerations as: handling aspects (e.g., weight and handles), indicators and scales, interchangeability, mechanical interference between accessories, alignment, interlocking, radiation, and mechanical safety (attachment security and collision avoidance), as well as positioning tolerances. The development of functional performance standards for medical linacs has served to focus attention on the need for standards for treatment
accessories.2l Recently, the IEC has developed standard terrninology for several accessories such as field blocks and tissue compensators, and is working toward a format specificationfor accessory holders, accessory rails, and shadow trays. As treatments become more complex and large accessory trays are attached to the radiation head, collisions between these accessories and the patient, couch, or couch attachments become more probable. The incident treatment beam should not intersect the couch accessory holders or rails, which may attenuate the beam by more than 40 percent. Such geometric problems can be expressed in terms of field sizes and gantry avoidance angles for specified treatment setups. These and other parameters have been incorporated into a brief computer program for planning treatments by Yorke.61 In the sections that follow, accessories are categorized by usage. They encompass both x-ray and electron therapy treatment modalities. References that pertain to several categories are listed under their primary usage. Goerl4 reviewed the role of treatment aids and accessories in radiation therapy treatment with emphasis on port films, electron arc therapy, patient immobilization and positioning, as well as record and verify systems. A comprehensive and well-illustrated exposition of radiotherapy accessories has been provided by hrdy.51 The recommendations and background considerations of a workshop on geometric accuracy and reproducibility in radiation therapy have been published.56 The 14recommendations of the workshop are divided into three categories: the significance of treatment accuracy, the achievement of accurate treatment, and the verification and evaluation of treatment accuracy. Boyers has reviewed patient positioning and immobilization devices, and their effect on radiotherapy treatment.
FIELD SHAPING SYSTEMS Most tumors are irregular in shape. Therefore, auxiliary absorbing shadow blocks are frequently used to modify the customary rectangular shaped x-ray fields provided by the conventional adjustable beam limiting collimator. Alternatively, a multileaf collimator having (e.g., multiple leaves of 1-cmeffective width or less at 100-cm SAD) can be used to approximate an irregularly shaped field. (see Chapter 2, pp41-43). Heavy metals such as tungsten, lead, or low melting point lead alloys (Lipowitz metals), such as Cerrobend, in lead equivalent thicknesses of 7.5 to 10 cm are needed to reduce the transmission of high energy x-rays to 5 percent or less for shadow blocks. Thicker shielding shadow blocks are often not warranted because of the limiting effects of penumbra and scatter from the irradiated volume into the shielded tissues. It is impractical to reduce transmission through a shielding block to 1 percent, for example, by increasing its thickness if scatter and penumbra already contribute 3 percent. The narrow beam transmission thicknesses of Lipowitz metal to the 1 percent level at 2,4,10, and 18 MV and for 60C0 has been reported.18 Lead or tungsten field blocks are frequently provided in an assortment of useful cross-sectional
TREATMENT ACCESSORIES
shapes (e.g., rectangles, triangles, and arcs of circles) for repetitive use. They are positioned on an accessory holder placed on the radiation head. Lipowitz metal blocks have become increasingly popular since they have a high physical density, a low melting point temperature, and can be cast in high density Styrofoam molds into irregular shapes to completely define an irregularly shaped field for an individual patient and provide for beam divergence.49 For supine orprone patients, a position used in the treatment of lymphomas, field blocks may be placed on a tray supported over the patient.45 Figure 12-4 shows the arrangement of shadow blocks employed for such treatments. Once a patient has completed treatment, these blocks are conveniently melted and recast into new shapes for patients treated at a later date. Several centers have described systems having automatic cutting machines for the molds and for casting and using such b1ocks.22.37.49.54 Thinner transmission blocks, which transmit 30-50 percent dose, are often used to reduce dose to sensitive treated tissues such as lung, while preserving the dose fractionation schedule.44.46 Their use insures an optimal radiobiological effect for such tissues as well as for nearby tissues irradiated to a larger dose. This regime is preferable radiobiologically to using 5 percent transmission blocks and removing them for one-half of the treatment days, for example. An ingenious, synchronous protection and field shaping system employing a curved absorber has been described for rotation therapy.5oUsing such a system, the curved spinal cord
205
of an individual can be protected by rotating an appropriately shaped shield, placed in the beam between the patient and x-ray target, so as to accurately shadow the continuously changing spine outline during rotation therapy. The plastic acrylic trays used to mount shielding blocks degrade the x-ray buildup region by contributing low energy electrons to the patient skin surface. The attendant, undesirable increase in skin dose can be reduced, by as much as 15 percent for large fields by substituting a clear leaded acrylic in the tray.4 Complex treatments, such as breast, entail the use of abutting fields with the attendant hazards of underdose and overdose associated with their divergence at the junction of the fields. Increasingly, such techniques incorporate field half-blocks in which one-half of the treatment field is blocked by placing one edge of the half-block along the beam central axis, thereby eliminating the divergence at the field edge near the ribs.35 The halfblock is by necessity, inconveniently large and heavy, and may restrict simultaneoususe of other accessories. Several treatment units now provide field half-blocking by the use of independently positioned beam limiting collimatorjaws.33Here, one or both pair of collimator jaws is constructed to provide either symmetric or asymmetric fields with respect to the central axis of the beam. For such fields, the asymmetric effects of flattening filter and scatter in the patient influence beam quality and the shape of isodose curves. The dosimetry of such asymmetric fields has been studied by Khan et al.29and Palta et a1.47 Electron fields are shaped with adjustable collimators or detachable applicators providing a small number of square or irregular field inserts. Thinner absorbers, typically of 1 to 2-cm equivalent lead thickness, can be used for irregularly shaped electron fields. They may be cast in thin, tailored Styrofoam molds as for high energy x-ray blocks. A technique for constructing square field or irregularly shaped field Cerrobend inserts has been described.52 The electron beam defining aperture is usually positioned at a small distance (e.g., 5 cm) from the incident skin surface. This separation allows for an unflat patient surface and permits observation of the light field on the patient's skin. The x-ray collimators provide "backup" peripheral shielding for electron treatment apertures and usually, are positioned automatically to a slightly larger (e.g., 5-cm extra added) field dimension.
WEDGE FILTERS AND TISSUE COMPENSATORS
-
FIGURE 12-4 Typical lymphoma treatment region showing the area covered by two adjoining fields, the mantle and inverted-Y fields, and the arrangement of shielding blocks needed.
Anatomical body surfaces are often nonplanar and irregular; an incident beam uniform in cross section would result in a nonuniform distribution of dose at tumor depth in a plane orthogonal to the beam axis. Wedge filters and tissue compensators are employed in such situations to modify the incident beam and to provide a more uniform dose distribution at the tumor depth. Wedge filters are usually in the form of 1-D absorbing wedges, which tilt the isodose distribution through a given angle (e.g., 30") at 10 cm depth. Pairs of wedges are often needed when adding fields at angles less than 180". Wedges are
206
CHAPTER 12. PATIENT SUPPORT ASSEMBLY AND TREATMENT ACCESSORIES
commonly used for two fields at 90' to each other but at larger and smaller angles as well. In one wedge accessory system, the wedges slide into a slot of the accessory holder placed close to the collimators at the front of the radiation head. A second adjacent slot is provided for attaching other accessories, such as shadow blocks. At the isocenter, wedged fields typically cover 15-20 cm in the wedged dimension and 3040-cm perpendicular to the wedged dimension. The wedge factor, that is, the fraction of incident dose transmitted on the central axis, will vary with beam energy, wedge material, and wedge angle. Representative wedge factor values at 10MV are about 0.90 at 15"and 0.50 at 60". The wedge becomes thicker and the wedge factor is decreased if the wedge is designed to cover a wider field in the wedged dimension. There is need for high machine dose rates to compensate for the marked attenuation of large angle, large field wedges. Wedges are electrically interlocked to insure complete insertion into the wedge slot and are labeled and individually coded to identify the particular wedge angle and direction of insertion, that is, 0" or 180" at the console. Miller and van de Geijn41 described a modification of the wedge fault logic to accommodate and code additional large field wedges in a Clinac 18/20. Most treatment units are provided with a selection of wedge filters that are mounted externally. A single 60" wedge, mounted inside the head and remotely positioned by motor control, may also be used as a universal wedge. Beams with effective wedge angles less than 60" are obtained by delivering an appropriate portion of the dose with the 60" wedged field and the remaining portion with an open (unwedged) field. Measurements of such a wedge by Petti and Siddon48 agree well with the model proposed by the manufacturer. The feasibility of providing wedge-shaped dose distributions by computered-controlled collimator jaw motion has been demonstrated by Kijewski et al.30 and has been called dynamic wedge. By adjusting the velocity of the jaw, one can produce fields with any wedge angle less than 60". The isodose curve in the principal plane can be made into a straight line of a predetermined angle at any desired depth and for an arbitrary field size.
The availability of independent jaw collimater systems and computer control has made the dynamic wedge an attractive option. As with the universal wedge", the technique simplifies wedge treatment setup since one no longer has to handle and install a physical wedge. Moreover, unlike physical wedges, the dynamic wedge delivers wedged-dose distributions that have a constant wedge angle over a larger portion of the radiation field with smaller effective perembra.32a Beam characteristics of wedged fields can be optimized independently for different wedge angles, field sizes, tissue depths, and x-ray beam energies. A dynamic wedge field may employ two adjacent segments having different wedge angles. A dual-pass technique, in which the dose is delivered by traveling the wedge field in both directions, provides isodose distributions adhering closely to the idealized prescription. A closed loop, servo-controlled computer system modulates the jaw speed and dose rate simultaneously to optimize and customize the dose distribution. An extention of this procedure will also incorporate the shape of the patient contour into the treatment planning system. Tissue compensators are employed for individual patients and provide 2-D dose modification by the use of shaped absorbers of wax, aluminum, brass, lead sheets, or lead alloys. The compensator is a 2D absorber constructed to compensate for missing tissue defined between a plane perpendicular to the beam axis and the patient's skin surface over the treatment area. Its affect is to provide a flat field over a plane perpendicular to the beam axis through the target volume at the tumor depth. Large field compensators do not compensate perfectly at other than one selected depth and can introduce hot spots at thinner body sections. The compensator is mounted on a plastic lamina that slides into an accessory slot. Placed between the target and skin surface, it takes the shape of a minified contour, which is the inverse of the incident skin surface (Fig. 12-5a). In one early compensator, aluminum columns of square I-cm cross section and various lengths were assembled into a 2D absorber matrix. It is convenient to vary the absorber length in multiples of 1-cm tissue equivalent,lo (see Fig. 12-5b).
FIGURE 12-5 . Tissue compensation. (a) The principle of shrinking the area of the tissue compensator, and (b)typical compensator constructed of aluminum square columns for 6 X 8-cm anterior field applied to the neck. (Courtesy of Ellis, from Ref. 10.)
207
TREATMENTACCESSORIES
An apparatus to measure the missing tissue depths at the patient by measuring the variation in distance of representative rays over a treatment field grid of 1-cm squares has been described.28 Another method involves plaster casts and photographs of successive contours, which are fabricated from lead sheets cut to the same shape with proper demagnification.31 In an alternative optical technique using lead sheet, a Moire' topographical photograph is obtained on the simulator.6 This is then employed to transform the depth of missing tissue contours in the photograph to a lead sheet compensating filter equivalent. One versatile and elaborate compensator system employs a contour milling machine design principle. In it, a stylus is moved over the treatment area while it simultaneously cuts out a 3-D Styrofoam mold using appropriate linkages. This mold is used to cast a Cen-obend compensator for the prescribed treatment geometry.l7,28,53 X-ray filters and compensators must be positioned at least 20 cm from the patient skin surface to limit the dose contribution from secondary electrons originating in them. Similarly, the use of lead acrylic accessory holder mounting plates limits secondary electron production and, hence, skin dose compared to materials such as Lucite.4 Filters designed to flatten the angular distribution of megavoltage x-ray beams are treated in Chapter 8, pages 147-148.
CT data for internal anatomical structures, as well as contours from a number of slices. Contour devices are frequently incorporated into simulators where they conveniently contribute to treatment planning information. One representative electromechanical contouring system is suspended above the patient on a telescopic support (see Figure 12-6).9 A system of two articulated arms and three precision potentiometers teminate in a probe arm contacting the patient's skin surface. The X-Y cartesian coordinates of the contour are obtained from the trigonometric based logic of the associated electronics. The contour coordinates for any number of cross sections are stored on tape and fed to the treatment planning computer. A similar system makes use of a graphics digitizer to provide computer compatible coordinates of patient contours.58 In another approach, an optical system, consisting of a light source, mirrors, and photomultiplier tube, is mounted on the simulator gantry.36 An associated feedback system and computer are used to calculate the distance from the isocenter of a slit of light reaching the patient in a sequence of angular positions of the gantry. The ultrasonic scanner has also been used for contour acquisition.
PATIENT CONTOUR SYSTEMS Patient contour systems are used to obtain a full scale, crosssection outline, usually a transverse section, of the patient's body surface for treatment planning. Frequently, this contour is limited to a central plane containing the beam axes, and usually containing the width or length of the field. Treatment planning is usually confined to this 2-D plane. However, several adjacent planes are needed for more complex treatments and a larger number when 3-D treatment planning is employed. Patient contours and the position of body organs can be obtained from CT scans, but substantial indexing and patient positioning problems are involved. Many patient contour devices are for use with the patient lying on the treatment couch in the prone or supine position. One of the earliest and simplest methods of contour acquisition is to shape one or more lengths of lead solder wire to approximate the patient contour and trace the outline(s) onto paper. This method and an alternative circular device, which incorporates a sequence of adjustable radial rods, are described by Lanzl et al.32 Contouring devices may be a simple mechanical rho-theta combination involving a variable length arm (rho), which can move through a variable angle (theta) about a fixed point of attachment. The sequence of rho-theta values traced out by the distal point of the arm can be translated tox-y coordinate points defining the patient contour. More sophisticated optical and x-ray systems are also employed. The latter may incorporate
Carriage -
Oigin of Coordinates
-
FIGURE 12-6 Schematic representationof patient contouring devices. Pi, Pz,and P3 are potentiometers that measure the vertical height of the carriage, the angle 0, and the angle 4, respectively. (Courtesy of Doolittle, from Ref. 9.)
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CHAPTER 12. PATIENT SUPPORT ASSEMBLY AND TREATMENT ACCESSORIES
PATIENT IMMOBILIZATION DEVICES Many radiation treatment procedures are predicated on spatial accuracies of 3 to 5 mm and some to 1 mrn.15,16,38 Patients are usually unable to maintain the position of the relevant body anatomy to this accuracy during treatment without auxiliary immobilization devices. Many treatment fields incorporate a I-cm border around the tumor to account for breathing displacements, assumed tumor spread, inaccuracies of CT, and so on. Hence, such devices are often used to simplify a procedure and insure accuracy on a daily basis during a course of therapy. Boyers has recently reviewed patient positioning and immobilization. He discusses the consequences of malpositioning and the need for frequent position verification. Immobilization devices are frequently used in conjunction with head and neck treatments, which often involve normal tissue structures in close proximity to the tumor, with an attendant need for spatial accuracy. They include bite blocks,24 as shown in Figure 12-7, molds, and other devices incorporating a wide variety of approaches. Such devices are constructed or adjusted for the individual patient and resewed for their exclusive use during a course of therapy. A reduction from 16 to 1 percent in significant localization error (a shift of 1 cm or more of the treatment portal in any direction compared to the original localization
film) was found in one head and neck study when immobilization was employed.39 Plaster casts were one of the earliest immobilization devices. They have largely been supplanted by thermal-setting or vacuumcast plastic lamina or meshes and more recently, by light-cast plastics,34 which set with the application of UV light.8,12,42,60Vacuum-activated sand blankets and tailored Styrofoam molds can be employed in a wide variety of procedures.42The vacuum-activated sand blanket is especially useful with infants and small children. Individualized plastic immobilization devices are used where precise alignment is important, often with tremorous patients. Treatment couches and chairs often incorporate arm supports, hand grips, and leg supports.
MECHANICAL AND OPTICAL POINTERS The positions of radiation treatment field axes on the patient are often established by the coincidence of skin marks placedon the patient with the central axes of the entrant and exit fields. Attachable front and back mechanical pointers, together with the image of an optical fiducial cross projected on the beam axis, were among the earliest devices employed to establish beam entrant and exit axes. Since such skin marks canmove with respect tointernal anatomy due to weight loss, hanging fatty tissues, and so on, distances from bony landmarks must be used to confirm the location of axes through a deep seated tumor. Direct use of skin marks for positioning tumors of the lower abdomen is not recommended, especially for obese patients. Isocentric treatment procedures, wherein the tumor is centered at isocenter, lend themselves to optical pointing methods. Here, auxiliary wall and ceiling mounted lasers project fiducial crosses or fan beams through the isocenter (see Figure 128)20,59 Their intercepts are marked on the patient's skin and provide a convenient reference for daily setup. Simulation procedures and the easy transfer of coordinate information are enhanced by providing a duplicate group of such lasers and a mechanical front pointer in the simulator facility. The use of CT images in providing anatomical information required for radiotherapy treatment planning is well established but the procedure is often indirect. Endo et al.11 describes an auxiliary optical system attached to the CT unit, which provides patient treatment beam positioning information and makes direct use of the CT images.
PATIENT POSITIONING AND MOTION DETECTION
FIGURE 12-7 . Rite-block mounted at end of treatnlent couch for supine patients. The patient's head rest in the plastic head i n ~ n ~ o b i l ~he.r , The bite-block assembly a and bracket c are readily detachable.
An early patient motion detection system incorporated a comparison of orthovoltage or megavoltage port films taken just before and just after treatment.' Although megavoltage film contrast is low, a dual exposure technique that superimposes an over-sized or maximum field exposure onto the smaller treat-
209
REFERENCES
Ceiling Laser
!
Accelerator Rotation
Side Lasers
I Angle Indicator
rce
Vertical Scale FIGURE 12-8 . Location of the iscocenter, gantry rotation axis, treatment couch, and positioning lasers on a &MeV linac.
ment portal exposure can provide useful anatomical information at lower megavoltage energies. Visualizing anatomical differences with film becomes increasingly difficult at 10 MV and above. Several high energy accelerators now incorporate a low energy (e.g., 6 MV) x-ray port filmmode. More complex and sophisticated motion detection procedures make use of a transmitted fluoroscopic image and video cancellation techniques by image subtraction, which portray differences in the observed image that change with time. An early study of continuous visual monitoring of the treatment beam relative to the anatomy and contrast media, such as air or mercury, involved a 2-MV Van de Graaff beam.' The ~ransmittedimage was viewed on a fluorescent screen with a television pickup system, amplified and reproduced on a monitor at the control console. It was possible to make minor position corrections during treatment in the case of rotation therapy. Kelsey et al.27 adapted a closed circuit television (CCTV) system for the measurement of patient motion during therapy. A preliminary study showed that patients remain within ? 2.5 mm of treatment positions during 90 percent of treatments. Patient repositioning and motion detection have also been studied using a video cancellation system.7 Here, the correct position is stored on a video disk and compared with the live image from a CCTV using a video subtraction technique. Repositioning and patient movements of 1 mm are detectable. The effect of position errors in the radiotherapy of head and
neck cancer has been assessed by Marks and Haus.39 Rabinowitz et al.55 examined the accuracy of radiation field alignment in clinical practices by comparing treatment-totreatment variations in anatomy with respect to the field by comparing sequential portal films. They found a normally distributed standard deviation of 3-mm independent of site and field shaping technique. Discrepancies between the portal and simulator films were considerably greater and depended on the site of treatment. The mean worst-case discrepancy averaged over all sites was 7.7 mm. This suggests that random uncertainties in treatment setup were not responsible for the discrepancies and they are more likely due to systematic errors in the treatment relative to the treatment simulation. An extensive description of portal imaging may be found in Chap. 13, pages 224-237.
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CHAPTER 12. PATIENT SUPPORT ASSEMBLY AND TREATMENT ACCESSORTES
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44. Orton CG, AM Herskovic, GA Ezzell, JT Spicka, T Vitalis: Transmission blocking: Clinical and biological rationales. Int J Radiat Oncol Biol Phys 11:2155-2158, 1985. 45. Page V, A Gardner, CJ Karzmark: Physical and dosmetric aspects of the radiotherapy of malignant lymphomas. Radiology.96:619-626, 1970. 46. Palos B, HS Kaplan, CJ Karzmark: The use of thin lung shields to deliver limited whole-lung irradiation during mantle-field treatment of Hodgkin's disease. Radiology 101:441442,1971. 47. Palta JR, MA Komanduri, N Suntharalingam: Dosimetric characteristics of a 6 MV photon beam from a linear accelerator with asymmetric collimator jaws. Int J Radiat Oncol Biol Phys 14:383-387, 1988. 48. Petti PL, RL Siddon: Effective wedge angles with a universal wedge. Phys Biol Med 30:985-991,1985. 49. Powers WE, JJ Kinzie, AJ Demidecki, JS Bradfield, A Feldman: A new system of field shaping for external-beam radiation therapy. Radiology 108:407411, 1973. 50. Priomos BS: Synchronous protection and field shaping in cyclotherapy. Radiology 77591-599, 1961. 5 1. Purdy, JA: Radiotherapy accessories. Encyclopedia of Medical Devices and Instrumentation, in John G. Webster, (Ed) New York, Wiley-interscience, 1988 IV pp 2466-2483. 52. Purdy JA: Electron beam simulation applicators. Med Phys 10:911-912, 1983.
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53. Purdy JA, SD Henderson, WB Harms: Compensating filter system design and use in radiotherapy. Proceedings of the Tenth Varian Users Meeting, Palm Springs, CA, April 15-17, 1984, pp 3 8 4 5 . 54. Purdy JA, A Sorenson,J Eisenhoffer: A combination Styrofoam cutter/block checker for fabricating field shaping blocks. Int J Radiat Oncol Biol Phys 2:1209-1211,1977. 55. Rabinowitz I, J Broomberg, M Goitein, K McCarthy, and J Leong: Accuracy of radiation field alignment in clinical practice. Inter J Radiol Oncol Biol Phys 11:1857-1867,1985. 56. Reinstein LE; Recommendationsof the workshop on geometric accuracy and reproducibility in radiation therapy. Med Dos 12:27-33, 1987. 57. Siddon RL, GL Tonnesen, GK Svensson: Three-field technique for breast treatment using a rotatable half-beam block. Int J Radiat Oncol Biol Phys 7:1473-1477, 1981. 58. Simpson LD, R Mohan: Computer compatible patient contour plotter. Med Phys 4:215-219, 1977. 59. Thomadsen BR: Principles in positioning cross-projecting lasers. Med Phys 8:375-377, 1981. 60. Verhey LJ, M Goitein, P McNulty, JE Munzenrider, HD Suit: Precise positioning of patients for radiation therapy. lnt J Radiat Oncol Biol Phys 8:289-294, 1982. 61. Yorke, ED: The geometry of avoiding beam intersections and blocking tray collisions. Med Phys 16:288-291, 1989.
Treatment Simulators, Treatment Planning and Portal Imaging
TREATMENT SIMULATORS "The simulator is and has been a radiotherapy aid for one-quarter of a century and yet, to many, its purpose and use are still undefined. The simulator is a different aid to each clinician who sets out to use it, and the use they make of it changes and grows with experience." This observation appears in the introduction to a report entitled: Treatment Simulators.l7 The BIR Supplement 2316 with this same title describes recent developments and perspectives concerning simulators that include imaging and computer aspects. These are excellent, general references on simulators which, in addition to design aspects, identify their value and use and indicate how they are incorporated into radiotherapy treatment planning facilities. Design considerations find their expression in performance specifications and the quality assessment and assurance of them.16.69.85 McCullough and Earlegs provide an extensive assessment of the latter aspects, often with somewhat tighter tolerances and from the perspective of American equipment and practice. Considerations in selecting related x-ray equipment, generator, tube, and image amplifier are included. Figure 13-1 illustrates a contemporary radiotherapy simulator providing both radiography and fluoroscopy. A quasisimulator alternative, which provides x-ray field verification, has been demonstrated by mounting a diagnostic x-ray source on the side of a linic.9 The results of diagnostic x-ray examinations have traditionally been incorporated into radiotherapy treatment planning, but a full awareness of the differing perspectives and needs of therapeutic and diagnostic radiology is a relatively recent perception. Conventional, diagnostic radiographic procedures play an important role in identifying and defining the nature and extent of disease. However, the difficulty of obtaining accurate, quantitative, reproducible, spatial information from radiographs using conventional diagnostic x-ray equipment largely precludes its use. This quantitative information is
not readily and accurately transferable from such equipment to the x-ray treatment unit nor is it verifiable.22 The radiotherapy treatment simulator was developed in response to these limitations. The CT images provide additional information but again accurate transfer of spatial data to the simulator may not be readily accomplished.22 Traditionally, diagnostic radiology has been concerned with qualitative, planar information obtained from views of anatomy positioned optimally for a particular diagnostic study. In contrast, radiotherapy planning requires quantitative, volumetric information obtained with the patient in a treatment position on a patient support assembly, usually a treatment couch. Accurate knowledge of field size and placement is crucial in radiotherapy, but of lesser interest in diagnostic radiology. Megavoltage port films, albeit of poor quality, provide information on the patient in a treatment position. The advent of CT and MRI, with their expanded tumor localization capabilities, has dramatized the differences in need and our perception of them. It has brought home two important tenets: (a) Ordinary diagnostic x-ray equipment and procedures are inadequate for obtaining treatment planning information. (b) Radiologists are needed who are aware of the special needs of radiotherapy treatment planning. As a result, there has been a dramatic increase in the development and use of simulators during the past decade, and diagnostic radiologists with an oncology interest have become an emerging subspecialty. As usually employed, the CT unit supplements the simulator and provides transverse (cross-section) images of the anatomy in which treatments are often planned. When the treatment plan is agreed upon, it is verified on the simulator using the treatment fields prescribed. Several developments have demonstrated that CT can be incorporated into simulators.45,72.97.99~111a~121~122 Alternatively, simulation can be incorporated into a CT scanner and integrated with software into a comprehensive treatment planning ~ystem.97,99J22~133 In one such system this includes interactive 3-D irregular field software and a software con-
214
CHAPTER 13. TREATMENT SIMULATORS, TREATMENT PLANNING AND PORTAL IMAGING
FIGURE 13-1
. Treatn~entsin~ulatorproviding radiography and fluroscopy. (Courtesy of Varian)
trolled laser located in the gantry for localization and transfer of treatment portal information to the patient's skin.45 Other contemporary developments pertaining to simulation are reviewed on pages 219-220. Data obtained from simulation is first used in planning the individual patient's treatment. Simulator originated data together with data from the treatment plan selected can be used for initiating a patient's daily therapy record used during treatment, for a record and verify function, and as the first step in establishing a more comprehensive computer based archival rec0rd.~0,69 New imaging methods, particularly CT and MRI, have added significant anatomic information for treatment planning. The simulator, however, remains central to the localization of the target volume and the verification of the planned treatment.
MECHANICAL FEATURES High energy megavoltage x-rays, with their penetrating and sharply delineated beams, characterize modern radiotherapy. Their ability to deliver a high and uniform dose to a deep seated tumor, while avoiding a nearby sensitive organ, is a hallmark. A radiotherapy simulator, which provides readily accessible and transferable quantitative tumor localization information, is essential to the planning process that exploits this capability. Where incorporated, CT and anatomical coordinate information must also be transferred." To simplify the transfer of information, a simulator should duplicate the treatment setting in significant detail, particularly the spatial, mechanical, radiation and optical aspects, and incorporate identical or similar accessories. The patient setup procedure for therapy is thereby
facilitated by a simple, 1:l transfer of information using high quality radiographs, guided by fluoroscopy if desired. Such an implementation will best fulfill the twin goals of simulation: (a) Provision, for plaizizirzg purposes, of accurate localization of the target voluine with respect to scirrounding anatonzy. (b) Verijkatioiz, after planniizg, of the prescribed treatnzentfield size andpositioniizg. These are the goals that motivate simulator design. The related procedures of treatment planning and portal imaging are treated on pages 220-237. Simulators are designed to simulate a number of different treatment units employing a wide variety of treatment techniques. Many simulator performance specifications can be described in terms of x-ray characteristics, mechanical parameters, and their tolerances. Stringent mechanical tolerances are needed to obtain the requisite spatial accuracy and to ensure reproducible setup of patients. Mechanical specifications such as field size range, SAD, SSD, isocenter height, and minimum and maximum couch top height, are important specifications. The simulator couch should duplicate the treatment couch as closely as possible, including removable panels and Mylar or "tennis racket" inserts, to simulate the sag of the patient on the treatment couch. Similarly, the accessories should be duplicates of those employed on the treatment unit. Accessory tray distance positioning on simulators must be compatible with the positioning on associated treatment units. Mechanical tolerances of simulators must be fully as strict (or possibly more strict) as those for treatment units as specified in the IEC standard.65.66 The lighter weight components of simulators can help ensure stricter mechanical tolerances so that simulation is unlikely to contribute significantly to overall error. How-
215
TREATMENT SIMULATORS
ever, the simulator must be designed to accurately position the heavy Cerrobend blocks used in treatment. Moreover, rigorous alignment 58 and quality assurance procedures '6,19.84-85 similar to those for linacs are needed. Often, tolerances are significantly more stringent than those for conventional diagnostic equipment from which they were developed. Several early simulators performed poorly through failure to meet these requirements. Mechanical and x-ray specifications suggested by a task group of the BIR may be found in Supplement 2316 and are reproduced in Table 13-1. This table is a summary of simulator specifications pertaining to: the gantry, the x-ray head and collimators, the x-ray tube and generator, the imaging device and the couch. These 1989 BIR recommendations were updated from an earlier 1981 tabulation.17 Some mechanical, as well as x-ray specifications, are discussed in more detail by McCullough, Blackwell, and Earle.84.85 The IEC standard conventions for linear and angular scale placement, direction of increasing values, and zero location for treatment units are recommended for simulators, as well as for treatment units.65 For convenience, both manual and motorized control of certain motions are suggested. The variable speed and powered motions specified in BIR Supplement 23 allow rapid, precise setting of these parameters. Digital readout of significant parameters is highly desirable. The BIR specifications, which involve anteroposterior (AP) and lateral exposures, are for a representative patient 25 X 45 cm in cross section with the image receptor 10-cm distal from the patient (see Figure 13-2). The design of the x-ray head and collimators (Table 13-1) is of major importance for simulation and treatment proce-
TABLE 13-1
- Summary of the simulator specifications
Height of isocentre above floor Angle of rotation at 5 100-cm SAD Angle of rotation at >100-cm SAD Isocentre accuracy, diameter Clearance between gantry and isocentre X-ray head and collimators Source-axis distance Beam limiting diaphragms at 100 cm Diaphragm rotation Beam delineating wires at 100 cm Source-skin distance indicator X-ray tube and generator Focal spot size Target angle Continuous rating of target Generator Radiographic output (minimum) Fluoroscopic output (minimum) Imaging device Film cassette and grid Image intensifier Scanning movements of the image intensifier Radial movement of the image intensifier Couch Couch top Rotation about couch support Rotation about isocentre Vertical movement Minimum couch height Longitudinal movement Lateral movement
5 1 1 5 cm >360°
0.03-1.0 rpm
+- 90"
2 mm ZllOcm 80-100 cma
0.5-5 cm/s
50 X 50 cm max > 220"
0.01 rpm
50 X 50 cm max 60-150 cm 0.3 X 0.3 mm 520" 500 HUIs 3 phase 500 mA, 90 kV 6 mA, 125 kV 2 3 5 cm sq 12 in.
Manual rotation
t 2 0 cm
3cds
-1Oto-60cm
3cds
220
X
45 cm
360" t 100 +2 to -50 cm 30 MeV Symmetry (Ratio of symmetrical points) Maximum ratio of absorbed dose (at D-max) 5X5to30X30cm to maximum square Wedge filtered x-ray fields Wedge factor Wedge angle Electron radiation Flatness (shape of isodose contours) 80% contour to geometric edge, at base depth 90% contour to geometric edgelcomer at SMD Maximum ratio of absorbed dose at D-max to absorbed dose on axis at SMD Stability of flatness with angular position of gantry and beam limiting system Symmetry (ratio of symmetrical points) Maximum ratio of absorbed dose at 0.5 mm depth to absorbed dose on axis at SMD Penumbra
(mfr) (mfr) (mfr) (mfr) +3%, + 3 mm* (mfr) (mfr) 20.1 cm 5 1.6 (mfr) +3%, 2 2 mm*
296
APPENDIX B. SURVEY OF MEDICAL LJNACS
TABLE B-4
-
(Continued) Indication of radiation fields X-radiation Numerical field indication (% is of field size) Greater than 20 cm X 20 cm to maximum square Light field indication: Edges (% is of field size): at normal treatment distance, 5 cm X 5 cm to 20 cm X 20 cm at 1.5 X normal treatment distance, 5 cm X 5 cm to 20 cm X 20 cm Center: NTDl 1.5 X NTD Reproducibility Numerical field, light field edge Electron radiation Numerical field indication Light field indication, edges Geometry of x-ray beam limiting systems Illuminance and penumbra of light field Average illuminance at NTD Edge contrast ratio Indication of radiation beam axis Entry, x-radiation (NTD 25 cm range) Entry, electron radiation (NTD ? 25 cm range) Exit, x-radiation (NTD to NTD 50 cm) Isocentre Displacement of x-ray beam axis Displacement of indication of isocentre Indication of distance along radiation beam axis from isocentre From radiation source Zero position of rotation scales Gantry, beam limiting device, table, table top Congruence of opposed radiation fields at isocentre Movements of the patient table Horizontal displacement for 20 cm vertical change Displacement of rotation axis from isocentre Angle between table and table top rotation axes Table height: 30 kg, retracted to 135 kg, extended Table top lateral tilt from horizontal Deviation of table top height with lateral displacement
3 mrn, 1.5%* 5 mm, 1.5%*
2 mm, 1%* 4mm, 2 %* 214 mm
2 rnm 2 mm + 0.5" 40 lux 400%
+
+
*=Whichever is greater NTD=normal treatment distance (100 cm SAD) SMD=standardmeasurement depth (10 cm, 90 cm SSD)
2 2 mrn +2 mm 2 2 mm +5 mm
Miscellaneous
C-1 ABBREVIATIONS ORGANIZATIONS AAPM ACR ANSI ASTRO BIR BRH CDRH CRCPD CROS CRP DHHS EPA ICRP ICRU IEC IEEE IPSM IS0 NBS NCRP NEMA NRC RPC SLAC SSRCR
American Association of Physicists in Medicine American College of Radiology American National Standards Institute American Society of Therapeutic Radiology and Oncology British Institute of Radiology Bureau of Radiological Health Center for Devices and Radiological Health Conference of Radiation Control Program Directors Committee on Radiation Oncology Studies Center for Radiological Physics Department of Health and Human Services Environmental Protection Agency International Commission for Radiation Protection International Commission on Radiation Units and Measurements International Electrotechnical Commission Institute of Electrical and Electronic Engineers Institute of Physical Sciences in Medicine International Organization for Standards National Bureau of Standards National Council on Radiation Protection and Measurements National Electrical Manufacturers Association National Regulatory Commission Radiological Physics Center Stanford Linear Accelerator Center Suggested State Regulations for Control of Radiation
TECHNICAL ACRONYMS ac AFC A1
Alternating current Automatic frequency control Artificial intelligence
AP BEV CART CCD CCTV CMOS CRT CT CW dc De-Q dm
Dm, DR DRR EMF EM1 ES fwhm EWTM GOS HBI HiNIL HVL IA IC I1 IORT LAN LAT MCII MLC MRI MTBF MTF MU NTP PA PACS PBI
Anterior posterior Beam's eye view Computer assisted'radiation therapy Charge coupled device Closed circuit television Complementary metallic oxide semiconductor Cathode ray tube (terminal) Computerized tomography Continuous wave Direct current Decreaser of Q Depth at dose maximum Dose at d, Digital radiography Digitally reconstructed radiograph Electromotive force Electromagnetic interference Expert system Full width half-maximum Full width tenth maximum Gadolinium oxysulfide Hemibody irradiation High noise immunity logic Half-value layer Image amplifier Integrated circuit Image intensifier Intraoperative radiation therapy Local area network Lateral Multichannel image intensifier Multileaf collimator Magnetic resonance imaging Mean time between failures Modulation transfer function Monitor units Normal temperature and pressure Posterior anterior Picture archival and communications system Partial body irradiation
APPENDIX C . MISCELLANEOUS
PC PET PFN PLO PRF PRR PSA PSIG QA QDE RAM rf ROM RPC RPM RTTP RV SAD SCR SIT SSD SW TBI TCP TE TEM TFD TLI TM TNI TSD TSET TTL TV TVD TVL TW VCO VDT VSWR
Printed circuit, personal computer Positron emission tomography Pulse forming network Phase locked oscillator Pulse repetition frequency Pulse repetition rate Patient support assembly Pounds per square inch guage Quality assurance Quantum detection efficiency Random access memory Radio frequency Read only memory Radiological physics center Revolutions per minute Radiotherapy treatment planning Record and verify Source axis distance Silicon controlled rectifier Silicon intensified target vidicon Source skin distance Standing wave Total body irradiation Tumor control probability Transverse electric Transverse electromagnetic Target film distance Total lymphoid irradiation Transverse magnetic Total nodal irradiation Target skin distance Total skin electron therapy Transistor transistor logic Television Tenth value distance Tenth value layer Traveling wave Voltage controlled oscillator Video display terminal Voltage standing-wave ratio
C-2 SYMBOLS A
8, B c C C C OC ~ G Y
Ampere Atomic weight Angstrom Magnetic flux density Velocity of light Capacitance Cosine-like trajectory Coulomb Degrees Celsius Radiation absorbed dose
d d D D e esu eV E
f
fc fc F F g &!
G G G GY h H H Hz I J J k keV kV ~ V P kW K OK 1 lm Ix L m m m m mA mAs M M MV MeV MIPS MVP MW n n n n N
Day Depth Absorbed dose Electric flux density Electron Electrostatic unit Electron volt Electric field intensity Frequency Cut-off frequency Foot candle Farad Force Gram Magnetic field gradient Admittance Gauss Giga (109) Gray Hour Henry Magnetic field intensity Hertz Electric current Current density Joule Kilo (103) Kiloelectron volt Kilovolt Kilovolt peak Kilowatt Coupling factor Degrees kelvin Length Lumen Lux Inductance Mass Meter Milli (10-3) Minute Milliampere Milliampere seconds Magnetic moment Mega (106) Megavolt Megaelectron volt Million instructions per second Megavolt peak Megawatt Nano (10-9) Neutron Magnetic field gradient index Turns ratio Newton
299
C-3 UNITS
Oe P P* P 4
Q ' 0
rad rem rms R R S
sr S S S T T T t
T
U v
v
W W
Oersted Momentum Momentum in units of Power Electric charge Quality factor of resonant circuit Range (continuous slowing down approximation) Radian Roentgen equivalent man (biological dose) Root mean square Resistance Roentgen Second Steradian Sine-like trajectory Surface area band 3000-MHz frequency region Kinetic energy Temperature Tesla Time Transit time factor Stored energy Velocity Volt Watt Total energy (rest plus kinetic) Rest energy Radiation length Atomic number Impedance (to ac or rf) One-dimensional Two-dimensional Three-dimensional
C-3 GREEK SYMBOLS TYPICAL USE
NAME a
Alpha
3 I
Beta
Gamma
Gamma Delta
Epsilon
Attenuation constant Alpha particle Velocity relative to light Beta particle (ray) Phase velocity Coupling factor Total energylrest energy Gamma ray Propagation constant Reflection coefficient Displacement Delta ray Skin depth Ernittance
Permittivity Dielectric constant Zeta Eta Theta Iota Kappa Lambda
Nu Xi Omicron Pi Rho
Sigma
Tau Upsilon Phi Chi Psi Omega Omega
Efficiency Instrinsic impedance of medium Angle Propagation wavenumber Wavelength Cut-off wavelength Guide wavelength Wavelength in free space Micro Permeability (magnetic) Micron
Circle circumferenceldiameter Radius of curvature Resistivity Reflection coefficient Standard deviation of normal distribution Conductivity Time constant Velocity Angle, phase Angle, phase Angular frequency (21~8in radians Impedance Resistance
C-4 UNITS In the portions of this book more related to engineering, such as Chaps. 2 to 5, 1 0 , 11, the rational system of units is used as introduced by Georgi. Lengths are in meters, mass in kilograms, time in seconds. Its primary advantage in relation to electricity and magnetism is that the units of all the primary electric quantities are those actually measured. Current is in amperes, potential in volts, impedance in ohms, power in watts. Force, in newtons, is the product of mass in kilograms and acceleration in meters per (second).2 1 newton = 105 dynes. Energy, in joules, is the product of force and distance. 1 joule = 1 newton meter = 107 ergs. The unit of charge is the Coulomb. Capacitance in farads = Ih (charge in Coulombs)2l(energy in joules). The dielectric constant of free space is eO =(!46~r) X 10-9 faradslmeter. Electric field is in volts per meter. Magnetic field intensity, H, from Ampere's law, has dimensions of current in amperes divided by length in meters.
300
APPENDIX C. MISCELLANEOUS
The magnetic flux density, B, is defined in webers, such that a rate of change in magnetic flux of one weber per second will generate an electromotive force of one volt. One weber is 108 maxwells or "lines" and one weber per square meter is 104 gauss. The permeability, p = B I H is IT X 10-7 henrylmeter. The henry is the unit of inductance, having units of volts induced by a rate of change of current in amperes per second. In some portions of this book, which are more related to physics, a rational system of centimeter-gram-second (cgs) practical units is used. Electromagnetic quantities are in volts, amperes, coulombs, ohms, and watts as in the meter-kilogramsecond (rnks) system, but length and mass are in centimeters and grams, hence involving conversion factors of 10-2 and 10-3 to convert to mks units.
C-5 TERMINOLOGY Absorbed dose: Mean energy imparted by ionizing radiation to matter. The special name of the unit of absorbed dose is the gray (Gy). It equals one on joule per kilogram (J. kg- I). Accelerator: Device for producing beams of high-energy electrons. The electron gun in a television tube is a type accelerator. Ampere: Aunit of electriccurrent. Since 1950,by international agreement, the ampere is defined in terms of the attractive force that occurs between two conductors carrying this current. The attractive force can be interpreted on the basis of magnetic forces--one conductor carrying a current generates a magnetic field at the other conductor, and the current flowing in this second conductor is then influenced by this magnetic field. The current flowing through a 100-watt bulb in an ordinary 110-volt house circuit is about 1 ampere. Attenuation: Reduction of a radiation quantity upon passage of the radiation through matter resulting from all types of interaction with this matter. The radiation quantity may be, for example, the particle flux density or the energy flux density. Note: Attenuation does not include the geometric reduction of the radiation quantity with distance from the radiation source. Bandwidth: 1. (continuous frequency band) The difference between the limiting frequencies. 2. (device) The range of frequencies within which performance, with respect to some characteristic, falls within specific limits. See: radio receiver. 3. (wave) The least frequency interval outside of which the power spectrum of a time-varying quantity is everywhere less than some specified fraction of its value at a reference frequency. Warning: This definition permits the spectrum to be less than the specified fraction within the interval. Note: Unless otherwise stated, the reference frequency is that at which the spectrum has its maximum value.
Beam dynamics: That branch of mechanics that deals with the motion or response of an electron beam under the influence of forces. Beam loading: The reduction of accelerated beam energy due to extraction of power from the accelerating electromagnetic field by the accelerated electron beam. Beta particle: An electron, either positively or negatively charged, emitted from a radioactive nucleus. Betatron: Electron accelerator in which an increasing magnetic field maintains a stable orbit and electrons are accelerated by an electric field produced by the increasing magnetic flux within the orbit. Build-up: In a material irradiated by a beam of x- or gammarays the increase in absorbed dose with depth below the surface is called the build-up. This is due to (a) the forward moving nature of the secondary electrons produced in the material, as well as (b) a build-up of scattered photons due to multiple scattering in broad beams of radiation. For high-energy beams process (a) is the more important. Capacitance (capacity): The property of a system of conductors and dielectrics that permits the storage of electrically separated charges when potential differences exist between the conductors. Centigray: 0.01 gray. 1 cGy equals one rad. (see gray) Choke joint : A connection designed for essentially complete transfer of power between two waveguides without metallic contact between the inner walls of the waveguides. It typically consists of one cover flange and one choke flange. Complex permeability: The complex (phasor) ratio of induction to magnetizing force. Notes: (1) This is related to a phenomenon wherein the induction is not in phase with the total magnetizing force. (2) In anisotropic media, complex permeability becomes a matrix. Conduction current: Through any surface, the integral of the normal component of the conduction current density over that surface. Note: (1) Conduction current is a scalar and hence has no direction. Coulomb: A unit of electrical charge. One coulomb of charge passing a section in a conductor every second is equivalent to a current of 1 ampere. Coupling: The circuit element or elements, or the network, that may be considered common to the input mesh and the output mesh and through which energy may be transferred from one to the other. Coupling aperture (coupling hole, coupling slot): An aperture in the bounding surface of a cavity resonator, waveguide, transmission line, or waveguide component which permits the flow of energy to or from an external circuit. Coupling capacitance: The association of two or more circuits with one another by means of capacitance mutual to the circuits.
C-3 UNITS
Coupling coefficient (coefficient of coupling): The ratio of impedance of the coupling to the square root of the product of the total impedances of similar elements in the two meshes. Notes: (1) Used only in the case of resistance, capacitance, self-inductance, and inductance coupling. (2) Unless otherwise specified, coefficient of coupling refers to inductance coupling, in which case it is equal to M/(LIL2)1/2, where M is the mutual inductance LI the total inductance of one mesh, and L2 the total inductance of the other. Coupling, inductance (interference terminology): The type of coupling in which the mechanism is mutual inductance between the interference induced in the signal system by a magnetic field produced by the interference source. Curie (Ci): The previously used special unit of activity equal to 2.7 X 1OlOdisintegrationsper second. 1 Ci = 3.7 X 1010 Bq. Delineator: Ameans for defining the border which outlines the simulated radiotherapy radiation field. Depth dose: Absorbed dose at a specified depth beneath the entrance surface of the irradiated object, usually on the radiation beam axis. Diaphragm: Beam limiting device with either a fixed or an adjustable aperture in practically one plane. Diode: A two-electrode electron tube containing an anode and a cathode. Dispersion (fiber optics): A term used to describe the chromatic or wavelength dependence of a parameter as opposed to the temporal dependence which is referred to as distortion. The term is used, for example, to describe the process by which an electromagnetic signal - is distorted because the various wavelength comvonents of that signal have different propagation characteristics.The term is also used to describe the relationship between refractive index and wavelength. Displacement current (any surface): The integral of the normal component of the displacement current density over that surface. Note: Displacement current is a scalar and hence has no direction. Displacement current density (any point in an electric field): The time rate of change in SI units (International System of Electrical Units) of the electric flux density vector at that point. Dispersion relation (radio wave propagation): In a sourcefree region, the functional relation between angular frequency w and wave vector k for plane waves with the exponential factor explj(wt-k-r)]. Dose monitor unit: In a dose monitoring system, arbitrary unit in which a quantity is displayed and from which absorbed dose can be calculated. Electric field (radio wave propagation):A state of the region in which stationary charged bodies are subject to forces by virtue of their charges.
-
-
301
Electric field: A vector field of electric field strength or of electric flux density. Electric flux density (electric displacement density) (electric induction): A quantity related to the charge displaced within the dielectric by application of an electric field. Notes: (1) Electric flux density at any point in an isotropic dielectric is a vector that has the same direction as the electric field strength and a magnitude equal to the product of the electric field strength and the absolute capacitivity. The electric flux density is that vector point function whose divergence is the charge density, and that is proportional to the electric field in region free of polarized matter. The electric flux density is given by
where D is the electric flux density, q , ~is the absolute capacitivity, and E is the electric field strength. (2) In a nonisotropic medium, E becomes a tensor represented by a matrix and D is not necessarily parallel to E. Electric focusing (microwave tubes): The combination of electric fields that acts upon the electron beam in addition to the forces derived from momentum and space charge. Electromagnetic waves: A wave characterized by variations of electric and magnetic fields. Note: Electromagnetic waves are known as radio waves, heat rays, light rays, etcetera, depending on the frequency. A disturbance that propagates outward from any electric charge that oscillates or is accelerated; far from the charge it consists of vibrating electric and magnetic fields that move at the speed of light and are at right angles to each other and to the direction of motion. Electron: 1. (noun) An elementary particle containing the smallest negative electric charge of 21.60219 X 10-19 C. Note: The mass of the electron 9.10956 X 10-31 kg is approximately equal to 111837 of the mass of the hydrogen atom. 2. (adjective) Operated by, containing, or producing electrons. Examples: Electron tube, electron emission, and electron gun. Electron accelerator, linear: See: linear electron accelerator. Electron charge: Charge is atomic in character; that is, there is a smallest amount below which charge may not be divided. The smallest charge e,; when negative, resides on certain elementary particles like the electron and antiproton, and when positive, resides on such particles as the positive electron (positron) and the proton. e = 1.6 X 10-20 coulomb. Electron emission: The liberation of electrons from an electrode into the surrounding space. Note: Quantitatively, it is the rate at which electrons are emitted from an electrode. Electron gun (electron tubes): An electrode structure that produces and may control, focus, deflect, and converge one or more electron beams.
302
APPENDIX C. MISCELLANEOUS
Emittance: A quantitative measurement of the nonparallelism of a beam (that is, low emittance signifies high parallelism). Energy: The capacity to do useful work. It may be kinetic energy, which is energy of motion, or potential energy, which is some potential form, such as gravitational,chemical, electrical, or atomic. As discussed in conservation of energy, mass is also convertible into energy. Fluorescence: Luminescence that occurs essentially only during excitation. Focal spot to skin distance: In radiotherapy, distance from the reference plane of an effective focal spot to the point at which the reference axis intersects with the entrance surface. Frequency (periodic function) (whereintime is the independent variable): The number of periods per unit time. Frequencyband: Acontinuous range of frequenciesextending between two limiting frequencies. Note: The term frequency band or band is also used in the sense of the term bandwidth. Full width at half maximum: For a bell shaped curve, distance parallel to the abscissa axis between the points where the ordinate has half of its maximum value. Full width at tenth maximum: For a bell shaped curve, distance parallel to the abscissa axis between the points where the ordinate has one tenth of its maximum value. Gantry: In equipment for radiotherapy, part of the equipment supporting and allowing possible movement of the radiation head. Geometric unsharpness: Unsharpness of the recorded image due to the combined optical effect of finite size of the radiation source and geometric separation of the anatomic area of interest. Gradient: The maximum rate of change of a parameter or characteristic in a given direction. Gray (Gy): International System (SI) unit for absorbed dose. One gray is equal to the energy imparted by ionizing radiation to a mass of material corresponding to one joule per kilogram; it is equal to 100 rad. Group velocity: 1. (radio wave propagation) Of a traveling wave, the velocity of propagation of the envelope, provided that this moves without significant change of shape. The magnitude of the group velocity is equal to the recip rocal of the rate of change of phase constant with angular frequency. Note: Group velocity differs in magnitude from phase velocity if the phase velocity varies with frequency, and differs in direction from phase velocity if the phase velocity varies with direction. 2. (traveling wave) The velocity of propagation of the envelope, provided that this moves without significant change of shape. Notes: (1) The magnitude of the group velocity is equal to the reciprocal of the change of phase constant with angular frequency. (2) Group velocity differs in magnitude from phase veloc-
ity if the phase velocity varies with frequency and differs in direction from phase velocity if the phase velocity varies with direction. 3. (waveguide) Of a traveling wave at a single frequency, and for a given mode, the velocity at which the energy is transported in the direction of propagation. Harmonic components (harmonics): The harmonic components of a Fourier Series are the terms C, sin (nx 0,). Note: For example, the component that has a frequency twice that of the fundamental (n, 2) is called the second harmonic. Image intensifier: An x-ray image receptor which increases the brightness of a fluoroscopic image by electronic amplification and image minification. Impedance: The total opposition that a circuit presents to an alternating current, equal to the ratio of the voltage to the current in complex notation. Note: The ratio Z is commonly expressed in terms of its orthogonal components, thus:
+
where Z, R, and X are respectively termed the impedance, resistance, and reactance, all being measured in ohms. In a simple circuit consisting of R, L, and C all in series, Z becomes
Z =R
+ j(oL - l/oC),
where o = 27rf and f is the frequency. Incident wave: 1. (radio wave propagation) In a medium of certain propagation characteristics, a wave which impinges on a discontinuity or a medium of different propagation characteristics. 2. (forward wave) (uniform guiding systems) A wave traveling along a waveguide or transmission line in a specified direction toward a discontinuity, terminal plane, or reference plane. See: reflected wave; waveguide. Inductive coupling (communication circuits): The association of two or more circuits with one another by means of inductance mutual to the circuits or the mutual inductance that associates the circuits. Insertion loss: (data transmission)Resulting from the insertion of a transducer in a transmission system, the ratio of (1) the power delivered to that part of the system following the transducer, to (2) the power delivered to that same part of the system after insertion of the transducer. Interlock: A device used to assure proper and safe use of a radiation installation by monitoring (usually by electrical devices) the status, presence or position of various associated devices such as source position, collimator opening, beam direction, door closure, filter presence and preventing the production or emission of radiation if the potential for an unsafe condition is detected. Interruption of irradiation: Stopping of irradiation and
303
C-3 UNITS
movements with the possibility of continuing without reselecting operating conditions, (that means return to the ready state). Ion: In a gas, a charged particle is often referred to as an ion (from the Greek word meaning wanderer) because it can move under the influence of an electric field. A negative ion may be an electron that has been freed from an atom or molecule in the gas; it may also be an electron that has become attached to a neutral atom or molecule. A positive ion is an atom or molecule that has lost one or more electrons. Ionization: Any process by which a neutral atom or molecule loses or gains electrons, thereby acquiring a net charge and becoming an ion. Ionization chamber: Ionization detector consisting of a charnber filled with a suitable gas, in which an electric field, insufficient to induce gas multiplication, is provided for the collection at the electrodes of charges associated with ions and the electrons produced in the sensitive volume of the detector by ionizing radiation. Iris: (waveguide technique) A metallic plate, usually of small thickness compared with the wavelength, perpendicular to the axis of a waveguide and partially blocking it. Notes: (1) An iris acts like a shunt element in a transmission line: it may be inductive, capacitive, or resonant. (2) When only a single mode can be supported an iris acts substantially as a shunt admittance. Isocentre: In radiological equipment with several modes of movement of the reference axis around a common centre, centre of the smallest sphere through which the radiation beam axis passes. Isodose curve: A line, usually in a plane, along which the absorbed dose is constant. Joule (J): The absolute meter-kilogram-second unit of work or energy equal to 107 ergs or approximately 0.7375 footpounds. One kilowatt-hour, which is the unit of electricity used by power companies, is equal to 3,600,000joules. A mass of 1 kilogram moving with a velocity of 1 meter per second has a kinetic energy of precisely .5 joule. Kinetic energy: The energy that a body possesses because of its motion; in classical mechanics, equal to one-half of the body's mass times the square of its speed. Klystron: An evacuated electron-beam tube in which an initial velocity modulation imparted to electrons in the beam results subsequently in density modulation of the beam; used as an amplifier in the microwave region or as an oscillator. Line spread function: In an imaging system, distribution of the intensity from a line source along a straight line in a specified image plane where the straight line is normal to the image of the line source. Loss: 1. (power) (A) Power expended without accomplishing useful work. Such loss is usually expressed in watts. (B)
(communications).The ratio of the signal power that could be delivered to the load under specified reference conditions to the signal power delivered to the load under actual operating conditions. Such loss is usually expressed in decibels. 2. (waveguide) The power reduction in a transmission path in the mode or modes under consideration. It is usually expressed as a positive ratio, in decibels. Loss tangent: The ratio of the imaginary part of the complex dielectric constant of a material to its real part.
Maximum dose depth: In a phantom whose surface is at a specified distance from the radiation source, depth on the beam axis at which the absorbed dose attains a maximum value with the specified irradiated field dimensions and beam energy. Modulation transfer function (MTF): A mathematical entity that expresses the relative response of an imaging system or system component to sinusoidal inputs as a function of varying spatial frequency, which is often expressed in line pairs per millimeter (lplmm). The reference value most commonly used is that for zero frequency. The MTF can be thought of as a measure of spatial resolution of the detector system. Fourier transform of the line spread function. For a symmetrical line spread function, the modulation transfer function is the normalized fourier transform using the equation: L(x) cos 2wx& M(v) =
j+--L(xk&
where v is the spatial frequency, L is the line spread function, and x is the abscissa.
Negative electricity: The sign of the electric charge may be either positive or negative (terms introduced by Benjamin Franklin) When glass is rubbed with silk, the charge retained on the glass is positive and on the silk is negative. Nominal energy: As a characteristic of medical electron accelerators, radiation energy describing: for x-radiation the energy of electrons in a monoenergetic beam equivalent to the actual energies of the energy spectrum of electrons in the radiation beam striking the target; for electron radiation the energy of electrons in a monoenergetic beam equivalent to the actual energies of the energy spectrum of electrons in the radiation beam at the phantom surface at the normal treatment distance. Oscillator: An electronic circuit or an enclosed metallic structure in which electric and magnetic fields vary periodically at a specific frequency. Particle accelerator: Equipment for accelerating charged particles such as electrons, proton, deuterons and alpha particles to kinetic energies higher than corresponding to the voltage applied. Thus-Electron accelerator. Linear
304
APPENDIX C. M
accelerator-particle accelerator in which charged particles are accelerated along a straight path. Particle fluence: At a given point of space, the number of dN particles incident during a given time interval of a suitably small sphere centered at that point divided by the crossdN sectional area of da of the sphere = da Peak pulse power, carrier-frequency: The power averaged over that carrier-frequency cycle that occurs at the maximum of the pulse of power (usually one half the maximum instantaneous power). Penumbra: In radiology, spatial region around the radiation beam where the value of radiation flux is between two specified or specific fractions of the value that is measured on the radiation beam axis, these two values being measured in a same cross-section. Note: The existence of such spatial regions can be due to one or more of the following phenomena: -extra-focal radiation, -scattered radiation, -absence of lateral electron equilibrium, -pair production, -geometry of the radiation source and of the beam limiting system. Percentage depth dose: The percentage depth dose in an irradiated body is the ratio (expressed as a percentage) of the absorbed dose, D,, at any depth x to the absorbed dose, Do,at a fixed reference point on the central ray. Percentage
+
0.x
depth dose = 100 X-. For x-radiation produced at up to Do 400 kV the reference point is at the surface. For x-radiation above 400 kV and gamma teletherapy the reference point is at the position of the peak absorbed dose. For movingfield therapy it is often convenient to take the centre of rotation as the reference point. Periodic electromagnetic wave (radio wave propagation): A wave in which the electric field vector is repeated in detail in either of two ways: (1) at a fixed point, after the lapse of a time known as the period, or (2) at a fixed time, after the addition of a distance known as the wavelength. Periodic function: Afunction that satisfiesf(x) =Ax + nk) for all x and for all integers n, k being a constant. For example, sin (x + a) = sin (x + a + 2n1~). Phantom: An object used to simulate the absorption and scatter characteristics of the patient's body for radiation measurement purposes. Phasing: Timing of a particle or a pulse with reference to either an oscillation or a circulation. Phase constant (waveguide): Of a traveling wave, the space rate of change of phase of a field component (or of the voltage or current) in the direction of propagation, in radians per unit length. Phase modulation (PM) (data transmission) (information theory): Angle modulation in which the angle of a carrier is caused to depart from its reference value by an amount proportional to the instantaneous value of the modulating
function. Notes: (1) A wave phase modulated by a given function can be regarded as a wave frequency modulated by the time derivative of that function. (2) Combinations of phase and frequency modulation are commonly referred to as frequency modulation. Phase velocity: 1. (fiber optics) For a particular mode, the ratio of the angular frequency to the phase constant. See: axial propagation constant; coherence time; group velocity. 2. (of a traveling plane wave at a single frequency) The velocity of an equiphase surface along the wave normal. See: radio wave propagation; waveguide. 3. (radio wave propagation) Of a traveling wave at a single frequency, the velocity of an equiphase surface along the wave normal. 4. (waveguide) Of a traveling wave at a given frequency, and for a given mode, the velocity of an equiphase surface in the direction of propagation. Photon: A massless particle; a "quantum" of electromagnetic energy. Synonymous with x-ray and gamma. Pixel: A two-dimensional picture element in the presented image. Polarity: The orientation of voltage between electrodes that determines the direction of current flow. Positron: A positively charged electron (see also electron charge). Potential energy (of a body or of a system of bodies, in a given configuration with respect to an arbitrarily chosen reference configuration): The work required to bring this system from an arbitrarily chosen reference configuration to the given configuration without change in other energy of the system. Potential gradient: A vector of which the direction is normal to the equipotential surface, in the direction of decreasing potential, and of which the magnitude gives the rate of variation of the potential. Practical range: For an electron beam, depth in a phantom at which the tangent to the steepest point, on the almost straight portion, of the depth versus absorbed dose curve, intersects the extrapolated tail of the curve.
Preparatory state: State of equipment for setting essential operating conditions, if in the standby state the setting of these conditions is not possible. Proton: Stable elementary particle having a positive charge of 1.60219 X 10-19 and a rest mass of 1.67261 X 10-27 kg. Quantum mottle: The variation in optical density, brightness, CT number, or other appropriate parameter in an image which results from the random spatial distribution of the x-ray or light quanta absorbed at the stage of the imaging chain containing the minimum information content. This state is known as the quantum sink. Rad: A superseded term that is an acronym for radiation absorbed dose. One rad is equal to 0.01 joules per kilogram, or 0.01 gray.
305
C-3 UNITS
Radiation (ionizing): Any electromagnetic or particulate radiation capable of producing ions, directly or indirectly, by interaction with matter. Leakage radiation-All radiation coming from within the source assembly except for the useful beam. (Note: Leakage radiation includes the portion of the radiation coming directly from the source and not absorbed by the source assembly, as well as the scattered radiation produced within the source assembly). Scattered radiation-Radiation that, during passage through matter is changed in direction. (It is usually accompanied by a decrease in energy.) Stray Radiation-The sum of leakage and scattered radiation. Useful beam,-The radiation which passes through the opening in the beam limiting device and which is used for imaging or treatment. Radiotherapy simulator: A device which uses x-ray equipment to physically simulate a therapeutic radiation beam so that the treatment volume to be irradiated during radiotherapy can be localized and the position and size of the therapeutic radiation field can be confirmed. Ready state: State of equipment, in which all conditions, such as carrying out of confirming operations and any other satisfaction of interlocks are prevailing so that the intended operation of such equipment can be initiated by a single action. Resolution: In the context of an image system, the output of which is finally viewed by the eye, it refers to the smallest size or highest spatial frequency of an object of given contrast that is just perceptible. The intrinsic resolution, or resolving power, of an imaging system is measured in line pairs per millimeter (lplmm), ordinarily using a resolving power target. The resolution actually achieved when imaging lower contrast objects is normally much less. Reflected wave: 1. (data transmission) When a wave in one medium is incident upon a discontinuity or a different medium, the reflected wave is the wave component that results in the first medium in addition to the incident wave. Note: The reflected wave includes both the reflected rays of geometrical optics and the diffracted wave. 2. (waveguide) At a transverse plane in a transmission line or waveguide, a wave returned from a reflecting discontinuity in a direction opposite to the incident wave. Reflection coefficient (waveguide): At a given frequency, at a given point, and for a given mode of propagation, the ratio of some quantity associated with the reflected wave to the corresponding quantity in the incident wave. Note: The reflection coefficient may be different for different associated quantities, and the chosen quantity must be specified. The voltage reflection coefficient is most commonly used and is defined as the ratio of the complex electrical field strength (or voltage) of the reflected wave to that of the incident wave. Examples of other quantities are power or current. Repetition rate: Repetition rate signifies broadly the number of repetitions per unit time.
Resonance charging (charging inductors) direct current: The charging of the capacitance (of a pulse-forming network) to the initial peak value of a voltage in an oscillatory series resistance- inductance-capacitance (RLC) circuit, when supplied by a direct voltage. Resonance frequency (resonant frequency): 1. (network). Any frequency at which resonance occurs. Note: For a given network, resonance frequencies may differ for different quantities, and almost always differ from a natural frequencies. For example, in a simple series resistance-inductance-capacitance circuit there is a resonance frequency for current, a different resonance frequency for capacitor voltage, and a natural frequency differing from each of these. See: network analysis. 2. (crystal unit) The frequency of a particular mode of vibration to which, discounting dissipation, the effective impedance the crystal unit is zero. See: crystal. Resonance frequency of charging (charging inductors): The frequency at which resonance occurs in the charging circuit of a pulse-forming network. Note: In this document, it will be assumed to be the frequency determined as follows:
where fo
= resonance frequency of
charging
Co = capacitance of pulse-forming network
L
= charging inductance.
Resonance mode (laser-maser): A natural oscillation in a resonator characterized by a distribution of fields which have the same harmonic time dependence throughout the resonator. Resonant mode: 1. (general) A component of the response of a linear device that is characterized by a certain field pattern, and that when not coupled to other modes is representable as a single-tuned circuit. Note: When modes are coupled together, the combined behavior is similar to that of the corresponding single-tuned circuits correspondingly coupled. See: waveguide. 2. (cylindrical cavities) When a metal cylinder is closed by two metal surfaces perpendicular to its axis a cylindrical cavity is formed. The resonant modes in this cavity are designated by adding a third subscript to indicate the number of half-waves along the axis of the cavity. When the cavity is a rectangular parallelepiped the axis of the cylinder from which the cavity is assumed to be made should be desjgnated since there are three possible cylinders out of which the parallele-piped may be made. Resonant wavelengths (cylindrical cavities): Those given by A, = l[(l/Xc)2 + (1/2c)2]1/2 where Xc is the cutoff wavelength for the transmission mode along the axis, I is the
306
APPENDIX C. MISCELLANEOUS
number of half-period variations of the field along the axis, and c is the axial length of the cavity. Resonator, waveguide (waveguide components): A waveguide or transmission line structure which can store oscillating electromagnetic energy for time periods that are long compared with the period of the resonant frequency, at or near the resonant frequency.
Rise time (industrial control): The time required for the output of a system (other than first-order) to make the change from a small specified percentage (often 5 or 10) of the steady-state increment to a large specified percentage (often 90 or 95), either before overshoot or in the absence of overshoot. Note: If the term is unqualified, response to a step change is understood: otherwise the pattern and magnitude of the stimulus should be specified. Root mean square value (high voltage testing): The root mean square value of an alternating voltage is the square root of the mean value of the square of the voltage values during a complete cycle. Root-mean-square value (periodic function) (effective value*): The square root of the average of the square of the value of the function taken throughout one period. Thus, if y is a periodic function o f t
where Y, is the root-mean-square value of y, a is any value of time, and T is the period. If a periodic function is represented by a Fourier series, then:
Root-sum-square: The square root of the sum of the squares. Note: Commonly used to express the total harmonic distortion. Scintillation: In radiology, luminescence of short duration (of the order of a few microseconds or less) caused by a directly or indirectly ionizing particle. Shunt: A device having appreciable resistance or impedance connected in parallel across other devices or apparatus, and diverting some (but not all) of the current from it. Appreciable voltage exists across the shunted device or apparatus and an appreciable current may exist in it. Shunting or discharge switch: A switch that serves to open or to close a shunting circuit around any piece of apparatus (except a resistor), such as a machine field, a machine armature, a capacitor, or a reactor.
Sievert (Sv): The special name for the SI unit of dose equivalent. One sievert equals one joule per kilogram. The previously used unit, rem, is being replaced by the sievert. One sievert is equal to 100 rem. Signal-to-noise ratio: For video cameras, the ratio of input signal to background interference. The greater the ratio, the clearer the image. Skin effect: Concentration of currents on the surface of conductors nearest to the electromagneticfield source producing them. Spectrum: 1. (data transmission) The distribution of the amplitude (and sometimes phase) of the components of a wave as a function of frequency. Spectrum is also used to signify a continuous range of frequencies, usually wide in extent, within which waves have some specified common characteristic. 2. (radiation) A distribution of the intensity of radiation as a function of energy or its equivalent electric analog (such as charge or voltage) at the output of a radiation detector. Stand-by state: State of an equipment which can be maintained for long periods and from which it is possible to move into rapid operation. For medical electron accelerators, state in which working levels of vacuum, temperature and other parameters are maintained but without the possibility of selecting the essential operating conditions. Standing wave linear accelerator: Linear accelerator in which radiofrequency energy is reflected at both ends of a tube in such a way that the particles are accelerated in a standing-wave electromagnetic field. Target volume: In radiotherapy, region of the patient containing those tissues which are to be irradiated with a specified absorbed dose according to a specified time-dose pattern. Termination of irradiation: Stopping of irradiation without the possibility of restarting without reselection of all o p erating conditions, (that means return to the preparatory state): when the preselected value of dose monitor units is reached, or when the preselected value of elapsed time is reached; or by deliberate manual actions; or by the operation of an interlock; or by preselected value of gantry angular position in moving beam radiotherapy. Thermionic cathode: A cathode in which electron emission is produced by heat. Pertaining to the emission of electrons as a result of heat. Transverse electric (TE,,,,,J resonant mode (cylindrical cavity): In a hollow metal cylinder closed by two plane metal surfaces perpendicular to its axis, the resonant mode whose transverse field pattern is similar to the TE,, wave in the corresponding cylindrical waveguide and for which p is the number of half-period field variations along the axis. Note: When the cavity is a rectangular parallelepiped, the axis of the cylinder from which the cavity is assumed to be made should be designated since there are three such axes possible. See: waveguide.
C-3 UNITS
Transverse-electric wave: 1. (radio wave propagation) An electromagnetic wave in which the electric field vector is everywhere perpendicular to the wave normal. 2. (TE wave) (general) In a homogeneous isotropic medium, an electromagnetic wave in which the electric field vector is everywhere perpendicular to the direction of propagation. See: waveguide. 3. (TE,, wave) (rectangular waveguide) (hollow rectangular metal cylinder) The transverse electric wave for which m is the number of half-period variations of the field along the x coordinate, which is assumed to coincide with the larger transverse dimension, and n is the number of half-period variations of the field along the y coordinate, which is assumed to coincide with the smaller transverse dimension. Note: The dominant wave in a rectangular waveguide is TE,,,: its electric lines are parallel to the shorter side. See: guided waves; waveguide. 4. (TE,, wave)(circular waveguide)(hollow circular metal cylinder) The transverse electric wave for which m is the number of axial planes along which the normal component of the electric vector vanishes, and n is the number of coaxial cylinders (including the boundary of the waveguide) along which the tangential component of the electric vector vanishes. Notes: (1) waves are circular electric waves of wave is the circular electric wave order n. The with the lowest cutoff frequency. (2) The TE,,, wave is the dominant wave. Its lines of electric force are approximately parallel to a diameter. Transverse electromagnetic (TEM) mode: 1. (fiber optics) A mode whose electric and magnetic field vectors are both normal to the direction of propagation. 2. (waveguide) A mode in which the lon~itudinalcomDonents of the electric and magnetic fields are everywhere zero.
m,, m,,
w
Transverse magnetic (TM) mode: A mode whose magnetic field vector is normal to the direction of propagation.Note: In a planar dielectric waveguide (as within an injection laser diode), the field direction is parallel to the core-cladding interface. In an optical waveguide, transverse electric (TE) and TM modes cornspond to meridional rays. Transverse magnetic (TM,,,,nd resonant mode (cylindrical cavity). In a hollow metal cylinder closed by two plane metal surfaces perpendicular to its axis, the resonant mode whose transverse field pattern is similar to the TM,, wave in the corresponding cylindrical waveguide and for which p is the number of half-period field variations along the axis. Note: When the cavity is arectangular parallelepiped, the axis of the cylinder from which the cavity is assumed to be made should be designated since there are three such axes possible. See: waveguide. Transverse-magnetic hybrid wave (radio wave propagation): An electromagnetic wave in which the magnetic field vector is linearly polarized normal to the plane of propagation and the electric field vector is elliptically polarized in this plane.
307
Transverse-magnetic wave: 1. (radio wave propagation) An electromagnetic wave in which the magnetic field vector is everywhere perpendicular to the wave normal. 2. (TM wave) (general) In a homogeneous isotropic medium, an electromagnetic wave in which the magnetic field vector is everywhere perpendicular to the direction of propagation. See: waveguide. 3. (TM,, wave)(circular waveguide)(hollow circular metal cylinder) The transverse magnetic wave for which m is the number of axial planes along which the normal component of the magnetic vector vanishes, and n is the number of coaxial cylinders to which the electric vector is normal. Note: TMo,, waves are wave is the circular magnetic waves of order n. The circular magnetic wave with the lowest cutoff frequency. See: guided wave; circular magnetic wave; waveguide. 4. (TM,, wave) (rectangular waveguide)(hollow rectangular metal cylinder). The transverse magnetic wave for which m is the number of half-period variations of the magnetic field along the longer transverse dimension, and n is the number of half-period variations of the magnetic field along the shorter transverse dimension. See: waveguide. Traveling wave: The resulting wave when the electric variation in a circuit takes the form of translation of energy along a conductor, such energy being always equally divided between current and potential forms. Traveling wave linear accelerator: Linear accelerator in which radio-frequency energy is applied at one end of a tube and is absorbed (or recirculated) at the other end in such a way that particles are accelerated in a traveling electromagnetic field. Treatment session: Fractionation: A session is a treatment or group of treatments delivered in one visit. Fractionation is the splitting of a dose into a number of short sessions given over a longer period than would be required if the dose were given continuously in one session at the same dose rate. A fraction is a single session in a fractionated treatment. Overall time is the total time elapsing from the beginning to the end of a session or a series of sessions if the treatment is fractionated. Treatment volume: In radiotherapy, region in the patient, to which the prescribed absorbed dose is delivered. Triode: A three-electrode electron tube containing an anode, cathode, and a control electrode. Volt (V): The unit of potential difference or electromotive force in the meter-kilogram-second system, equal to the potential difference between two points for which 1 coulomb of electricity will do 1 joule of work in going from one point to the other. A battery of cells develops an electric-potential difference across its terminals by means of chemical-potential energy. The potential difference of an ordinary flashlight cell is approximately 1.6 volts.
308
APPENDIX C. MISCELLANEOUS
Waveguide: A system of material boundaries or structures for guiding electromagnetic waves. Usually such a system is used for guiding waves in other than TEM modes. Often, and originally, a hollow metal pipe for guiding electromagnetic waves. Wavelength: The distance, for example, between two successive crests of a wave. Many illustrations of waves are available, such as: mechanical waves on the surface of water, along strings, and in air, electromagnetic waves as in light and x-ray; and so-called matter waves of nuclear and atomic particles. Matter waves were predicted by de Broglie in 1923, and the matter waves of electrons were discovered by Davisson and by G.P. Thomson in 1925. These developments culminated in the theory of wave mechanics.
Wedge angle: Angle between the perpendicular to the radiation beam axis and the straight line passing through two points equidistant from the radiation beam axis located on the isodose curve, at the measurement depth specified in a phantom, and separated by a distance equal to half the irradiation field, for a given irradiation field dimension and beam energy. Wedge factor: For a given beam energy and irradiation field, ratio of the absorbed dose rate in a phantom on the radiation beam axis at the standard measurement depth with the wedge filter in place to the absorbed dose rate without the wedge filter. Wedge filter: Filter which effects stepless change in transmission over all or a part of the radiation field.
Index
Abbreviations, 297-298 Aberrations, second order, 131 Accelerator cavities, design of, 86-87 Accelerator control, 169- 188 computer control, 170 computer integration of radiotherapy, 181- 187 control console, 169, 177- 178 extreme dose, protection against, 176-177 interlock system, 169- 170,173- 176,256-257 miniaturization and, 170- 171 motion control system, 178- 180 operational states, 173 patient record keeping, 181 record and verify system, 180- 181 semiconductor devices, 171- 173 Accelerator facilities, 24 1-259 human engineering aspects, 257-258 megavoltage therapy accelerator facilities, 244-252 entry doors/mazes, 25 1 multimodality therapy installation, 244-247 patient obse~ation/communication,25 1-252 radioactivation of patient, 252 radioactive and toxic gas production, 252 shielding barrier design, 247-25 1 treatment room design, 247 planningloperational resources, 24 1-244 safety, 256-257 See also Accelerator maintenance Accelerator guide, multi-x-ray energy accelerators, 191- 193 Accelerator maintenance, 252-256 conventional maintenance, 252-253 downtime experience, 256 expert systems, 254 functional performance, periodic tests of, 255-256 test equipment/instrumentation, 254-255 usageldowntime, 256 Accelerator operational states, 170 Accelerator structures, 49,67-87 accelerator cavities, design of, 86-87 electron interaction with microwave field, 68-70
standing-wave accelerators, 76-82 traveling-wave accelerators, 70-76 standing-wave accelerators compared to, 82-86 Acceptance, beam transport, 115 Accessories, S d r e a t m e n t accessories, simulation accessories Achromatic bend magnet systems, 129-134 asymmetric 112 112-degreethree-sector uniform pole gap, 133 symmetrical 180" four-sector uniform pole gap (isochronous), 133- 134 symmetrical 270" single-sector hyperbolic pole gap, 129 symmetrical 270" single-sector locally tilted pole gap, 129 symmetrical 270" single-sector stepped pole gap, 129- 130 symmetrical 270" three-sector uniform pole gap: one Cx cross-over, 131- 133 two Cx cross-overs, 130- 131 Achromatic, singly, doubly, 129 Anatomical landmarks, 224 Applicator, See Treatment accessories Arc therapy, 41 Arteriovenous malformation (AVM), See Sterotactic surgery Automatic frequency control (AFC), 19,27, 102- 104 high-energy (klystron), 103- 104 low-energy (magnetron), 102- 103 Automatic wedge, 228 Auxiliary systems: gas dielectric system, 111- 113 pneumatic system, 113 vacuum systems, 107- 110 water cooling system, 110- 111
Beam collimators, 141, See also Collimation symmetric vs. asymmetric, 141 Beam emittance, 115 Beam loading: multi-x-ray energy accelerators, 197 standing-wave accelerators, 80-82
310
traveling-wave accelerators, 75-76 Beam mode, user preferences for, 3-4 electrons, 4 X-ray energies, 3-4 Beam optics of magnet systems, 115-136 achromatic bend magnet systems, 129-134 beam emittance, 122- 124 bent beam linacs, 115 electron motion in magnetic fields, 116- 122 isocenter height, effect of system choice on, 115-16 linac beam characteristics, effect on, 115 nonachromatic bend magnet systems, 125-129 straight-ahead linacs, 115 Beam stabilization, See Treatment beam stabilization Beam steering, 157-159, See also Field uniformity control coils, 122 Beam transport, 122 system terminology, 123 Beams eye view (BEV), 43,222-223 Beamstopper, 140 Beam waist, 124 Bent beam linacs: beam design, 115 standing-wave, 12- 13 vs. straight thru, 139 Bend magnet, 18 Betatron, 6-7,271-273 Blocking, See Treatment accessories, field shaping systems Bremsstrahlung (braking radiation), 275
Cancer incidence, 1 Cell repair, regeneration, redistribution, reoxygenation, 3 Chromaticity, 122 Circular microtron, 14,26 1-263 beam current/focusing, 263 cavity power requirements, 261-262 gantry, 263 injection methods for increased energy per orbit, 262 machines for radiotherapy, 263 magnet size, 262 phase stability, 262-263 Circulators, 95-97 Collimation, collimators, 41, 141, See also Multileaf collimators Compensators, See Tissue compensators Compton interaction, 279 Computer control, of accelerator, 170 of treatment, 41-46 Computer integration of radiotherapy, 181- 187 Conformal therapy, multileaf collimators and, 26,28-29, 4 1-43 Contamination of radiation beam, 150 Contour systems, 207 Contrast sensitivity, 225
INDEX
Control console, 177- 178 Cosine-like rays, 129 Coupling, 59-62 directional couplers, 98 Curative intent, 1, 15, 20
Defocusing, of electron rays, 119 Depth dose, 3 Distribution, 25 Stability, 25 Digital spatial filtering, 231 Diode guns, 191 Dipole magnetic field, electron motion in, 116- 117 Direct accelerators, 6 Directional couplers, 98 Dispersion, 62-63 Displacement, of electron ray, 115 Divergence, of electron ray, 115 Dose buildup, 3,23 Dose, depth, 33,37, See also Isodose, skin dose Dose monitoring, 157-168 electromagnetic interference and, 166- 167 field uniformity control, 161- 162 integrated dose/dose rate, 160- 161 ionization chamber, 157- 160, See also Electromagnetic interference multimodality treatment units, control of, 162-164 safety, 160-164 system, 157-162 treatment beam stabilization, 164- 166 Dual x-ray energy standing-wave linacs, 13 Dynamic therapy, multileaf collimators and, 41-43
Electromagnetic interference (EMI), 166- 167, 171- 173 with ionization chamber, MRI, pacemaker semiconductor devices and, 171- 173 Electron accelerators: history of, 6-13 betatrons, 6-7,271-273 direct accelerators, 6,267 microwave electron linear accelerators (linacs), 7- 13 recirculating electron accelerator, 13-15,267-268 resonant transformer, 270 synchrotron, 273 transformer-rectifier,268-270 Van de Graaff, 270-27 1 Electron beam generation, 278-279. See also Electron therapy scattering foils, 143- 144,279 spurious sources, 278-279 Electron boast fields, 189 Electron-positron pair production, 279 Electron guns: cathode, 67-68,191
design of, 68 diode, 191 multi-x-ray energy accelerators, 191 triode, 191 Electronic portal imaging, 227-238 lens to TV camera, 233-234 mechanically rotated multichannel ionization chamber, 232 mounting detector on linac, 235 multiwire sequentially pulsed liquid ionization chamber, 230-232 one-dimensional vs. two-dimensional detectors, 228-230 silicon detectors, two-dimensional array of, 234 silicon diode linear array, 230 tapered fiber optics to TV camera, 232-233 two-dimensional amorphous silicon array, 234-235 value of, 227-228 Electrons, 4 interaction with microwave field, 68-70 motion of electrons, 68-70 motion in magnetic fields, 116- 122 beam steering coils, 122 beam transport, 122 in dipole magnetic field, 116- 117 electron momentum, 116 in fringe field at dipole magnet edge, 117- 119 in quadruple magnetic field, 119- 120 in solenoid magnetic field, 120- 122 space harmonics, 70 Electron synchrotron, betatron and, 27 1-273 Electron therapy, 142- 145 beam current requirements, 163 electron scanning system, 144- 145 electron scattering system, 144 megavoltage, 37-38 microtrons vs. linacs for, 145 total skin, 38-39 Elementary linac, 50 Emittance, beam, 115 Energy, electron, specification and measurement, 143 Energy slit, 125, 190 Expert systems, accelerator maintenance and, 254 Extreme dose, protection against, 176-177
Facilities, See Accelerator facilities Field clamps, 119 Field light and rangefinder, 142 Field shaping systems, 204-205, See also Collimation, Multileaf collimator Field uniformity control, 161-162, See also Beam steering Filling time, 20 Filter, flattening, 141, 147- 148 Flattening filter, 18 Flexible waveguides, 97-98
Fluorescent screen, 225 Focusing, of electron rays, 119 Fractionation, of treatment course, 2 Frequency instabilities, 26 Fringe field at dipole magnet edge, electron motion in, 117-119 Functional performance, periodic tests of, 255-256
Gas dielectric system, 111- 113 Geometric image unsharpness, 225 Greek symbols, 299 Group velocity, 59 Guide, See Accelerator structures, waveguide structure, 11,71, 191-193 Guns, See Electron guns
Half-value-layer, 34 Head leakage radiation, See Radiation shielding Hemibody x-ray therapy, See Total-body/hemibody x-ray therapy High-energy (klystron) automatic frequency control, 103-104 H-plane tee, 98-99 Hybrid tees, 98-99
Illuminance (foot-candles), 233 Image plane, 122 Imaging smoothing, 23 1 Immobilization devices, 208 Impedance matching, 54-55 voltage standing-wave ratio and, 55 Initial seconds of each portal treatment, 26, 190 In-line standing wave linacs, 11- 12 Integrated dose/dose rate, monitoring, 160- 161 Interlocks, 169-170, 173-176,256-257 Interrupting radiation, 169- 170 Intraoperative radiation therapy, 39-40 Ionization chamber, See also Electromagnetic interference dose monitoring, 157- 160 mechanically rotated multichannel ionization chamber, 232 multiwire sequentially pulsed liquid ionization chamber, 230-232 Isocenter height, effect of magnet system choice on, 115- 116 Isocentric linac, 10 Isochronous, 133 Isodose, 33-34,37
Klystrons, 8,91-92 automatic frequency control, 103- 104 magnetrons vs., 191 multimegawatt klystron, 9
312
INDEX
Leakage radiation, See Radiation shielding L i c generated EMI, 171- 172 Local area network, 182 Localization radiograph, 224 Low-energy (magnetron) automatic frequency control, 102-103
Machine interlocks, 169,257 Machine performance requirements, treatment course fractionation and, 2-3 Magna-field therapy, See Total-body/hemibody x-ray therapy Magnetic resonance imaging, See MRI Magnetrons, 8,89-91 automatic frequency control, 102-103 klystrons vs., 191 Magnet systems: beam optics of, 115-136 isocenter height, effect of system choice on, 115- 116 Magnetic forces, 115 Magnetic minor, 129 Magnification, 119 Maintenance, See Accelerator maintenance Mean time between failures, 26 Mechanically rotated multichannel ionization chamber, 232 Mechanical pointers, 208 Medical electron accelerators: accelerator control, 169- 188 accelerator facilities, 241-259 beam mode, user preferences for, 3-4 betatron, 6-7,271-273 clinical requirements, 22-23,142 precise delivered dose at depth, 22 precise position/orientation/sizeof treatment fields, 22-23 wide variety of radiation modalities, 23 design challenges, 24-26 compactness, 24 dose precision, 24 energy stability, 25 high-dose rate with large fields, 24 initial seconds, 26 treatment beam stability, 24-25 uniform electron treatment beams, 25 uniform x-ray treatment beams, 25 design criteria, 20-2 1 electron accelerators, history of, 6-15,267 electron synchrotron, 27 1-273 energy conversion, summary of steps in, 20 equipment development, future direction of, 27-29 machine performance requirements, treatment course fractionation and, 2-3 major subsystems/components, 18-20,26-27 modulator/high-voltagepulse transformer, 18- 19
radio frequency power sourcelradio frequency power control, 19-20 microwave acceleration, principle of, 16- 18 need for, 1 reflexotron, 15,267-268 resonant transformer, 270 safety interlocking, 169- 188, See also Interlocking, Safety interlocking technology changes, 26 transformer-rectifierunits, 268-270 Van de Graaff generator, 270-271 Medical linacs, elementary description of, 15-16 example, 36-38 Medical microtron accelerators, 261-266 circular microtron, 26 1-263 beam current/focusing, 263 cavity power requirements, 261-262 gantry, 263 injection methods for increased energy per orbit, 262 machines for radiotherapy, 263 magnet size, 262 phase stability, 262-263 racetrack microtron, 263-266 accelerator structure power, 264 alignment precision, 265 configuration, 263-264 extraction, 265 injection, 264 machine for radiotherapy, 266 Megavoltage electron therapy, 37-38 beam current requirements, 163 Megavoltage therapy accelerator facilities, 244-252 entry doorslmazes, 25 1 multimodality therapy installation, 244-247 patient observation/communication, 25 1-252 radioactivation of patient, 252 radioactive and toxic gas production, 252 shielding barrier design, 247-25 1 treatment room design, 247 Megavoltage x-ray therapy, 34-35 beam current requirements, 163 Microtrons, See Medical microtron accelerators Microwave acceleration, principle of, 16- 18 Microwave cavities, 7-8 Microwave electron linear accelerators (linacs), 7-13,267 accelerator structures, 67-87 beam characteristics, 115 bent beam standing-wave linacs, 12-13 contemporary radiotherapy accelerators, 287-296 dual x-ray energy standing-wave linacs, 13,293-294 first orientable linacs for radiotherapy, 9- 11 first stationary linac for radiotherapy, 9 in-line standing wave linacs for radiotherapy, 11- 12 klystron invention, 8 magnetron invention, 8
313
INDEX
manufacturers' types, 287-296 microwave cavities, 7-8 microwave linac invention, 8-9 microwave principles for, 49-66 multimegawatt klystron invention, 9 operating parameters, 38, 163 pioneers, 13 standiig-wave accelerator guide, 11 Microwave energy switch, 192, 194 Microwave linac, 8-9 Microwave power sources/systems, 89- 104 automatic frequency control, 102- 104 circulators, 95-97 directional couplers, 98 flexible waveguides, 97-98 klystrons, 9 1-92 magnetrons, 89-91 radio frequency drivers, 92-95 rotary joints, 99 shuntJserieslhybrid tees, 98-99 water loads, 101- 102 waveguide bends and twists, 97-98 waveguide windows, 100- 101 Microwaves, 50-5 1 Miniaturization, accelerator control and, 170- 171 Modes, microwave resonator, 62-63 Modulation transfer function (MTF), 225,228-229 Modulator, high voltage, 16 Momentum trajectories, 131 Momentum, electron, 116 Motion control system accelerator, 178- 180 MRI (magnetic resonance imaging) EM1 interference with, 166 Multileaf collimator (MLC), 28,227-228 Multileaf collimators, dynamic and conformal therapy and, 41-43 Multimegawatt klystron, 9 Multimodality treatment units: control of, 162- 164 installation, 244-247 Multiwire sequentially pulsed liquid ionization chamber, 230-232 Multi-x-ray energy accelerators, 189- 199 design alternatives, 191- 199 accelerator guide, 191- 193 beam loading, 197 electron gun, 191 klystron vs. magnetron, 191 design challenges, 189- 191 clinical need, 189 dose distribution/calibration in initial seconds, 190-191 electron beam during acceleration, 190 energy stability, 190 performance requirements, 189- 190 non-contact-type side cavity energy switch, 197
standing-wave guide, switching from high to low x-ray energy in, 193-196 system feedback control philosophy, 199 traveling-wave guide, switching from high to low energy in, 193
Neutron activation and leakage, 150- 151 Nonachromatic bend magnet systems, 125-129 Non-contact-type side cavity energy switch, 197
Object plane, 122 Off-axis portal x-ray tube, 236 On-axis portal x-ray tube, 236-237 Operational states, accelerator, 170 Optical pointers, 208 Orientable linacs, 9- 11 Orthovoltage x-ray therapy, 33-34
Pacemaker, interference with, 166- 167 Palliation, 1 Patient: contour systems, 207 immobilization devices, 208 motion unsharpness, 226 observation/communication, 25 1-252 position/motion detection, 208-209 radiation interactions in, 279-281 radioactivation of, 252 record keeping, 181 Patient support assembly, 201-204 table support types, 201-203 table top, 203 treatment chair, 201,203-204 See also Treatment accessories Penumbra, 3 Periodic structures, microwave, 59-62 Personnel interlocks, 169 Phase space admittance, 264 Phase stability principle, 261 Phase space volume, 123 Phase velocity, microwave, 59 Photoelectric effect, 279 Photon intensity, angular distribution of, 275-276 Photon spectrum in portal imaging, 235-237 image contrast, dependence on x-ray energy, 236 Picture archival and communication system (PACS), 182 Pneumatic system, 113 Portal imaging: electronic, 227-238 off-axis portal x-ray tube, 236 on-axis portal x-ray tube, 236-237 photon spectrum in, 235-237
Portal imaging (cont.): radiographic (film), 224-227 PositionJmotion detection, 208-209 Pulse-forming network (PFN), 19 Pulse modulators, 105- 107 pulse-forming network (PFW), 105, 106
Quadruple magnetic field, electron motion in, 119-120 Quadruple vs. four pole magnet, 145 Quality of life, 1 Quantum detection efficiency (QDE), 225
Racetrack microtron, 14- 15,263-266 accelerator structure power, 264 alignment precision, 265 configuration, 263-264 extraction, 265 injection, 264 machine for radiotherapy, 266 Radiation beam: contamination of, 150 energy designation, 33 generation, See Treatment beam generation penetration, See Dose, depth types of, 33 Radiation head, geometric restrictions of, 138- 139 accessibility for service, 138 Radiation length, 276 Radiation shielding, 140- 141 materials, 140 neutron, 140 Radioactivation, 150- 151 of patient, 252 Radioactive and toxic gas production, 252 Radio frequency drivers, 92-95 Radiographic (film) portal imaging, 224-227 enhancement techniques, 226-227 physics of, 224-226 Radiosurgery, See Stereotactic radiosurgery Radiotherapy: computer integration of, 181- 185 goals of, 1-2 modalities, 33-48 arc therapy, 4 1 conformal therapy, 41-43 dynamic therapy, 41-43 intraoperative radiation therapy, 39-40 megavoltage electron therapy, 37-38 megavoltage x-ray therapy, 34-35 orthovoltage x-ray therapy, 33-34 rotational therapy, See Arc therapy stereotactic radiosurgery, 43-45 total-body and hemibody x-ray therapy, 35-37
total skin electron therapy, 38-39 organization, procedures, staffing, 2 18,221 Radiotherapy accelerator facilities, See Accelerator facilities Radius of curvature, 116 Recirculating electron accelerators, 13- 15 circular orbit microtron, 14 racetrack microtron, 14- 15 Reflexotron, 15,267-268 Record and verify system, 180- 181 Record keeping, patient, 181 Reflexotron, 15,267-268 Resonance/resonant cavities, 55-59 Resonant transformer, 270 Rotary joints, microwave, 99
Safety, accelerator facilities, 256-257 collision avoidance, 137 dose monitoring, 160- 164 electromagnetic interference, 166- 167 extreme dose, protection against, 176 interlocking, See Interlocks Scanned beam dosimetry, 150 x-ray therapy, 150 Scanning system: electrons, 42-43, 144- 145 x-ray, 42-43, 148-150 Scattering foils, electron, 279 Scattering system, electrons, 144 Series tees, 98-99 Shadow blocking, 224, See also Treatment accessories, field shaping systems Shielding barrier design, megavoltage therapy accelerator facilities, 247-251, See also 140- 141 Shunt impedance, 64-66 Shunt tees, 98-99 Side coupled SW accelerator structure, 7, 11 Signal to noise ratio, 231 Simulators, treatment, See Treatment simulators Simulator verification, treatment planning and, 221 Sine-like rays, 129 Skin dose, 34 Solenoid magnetic field, electron motion in, 120-122 Space harmonics, electrons, 70 Spatial resolution (line spread function), 231 Standing-wave accelerator guide, 11 Standing-wave accelerators, 76-82 beam loading and load line, 80-82 electron injection and bunching, 79-80 operation, theory of, 76-78 structures, 78-79 traveling-wave accelerators compared to, 82-86 See also Traveling-wave accelerators Stationary linac, 9 Stereotactic radiosurgery, 43-45
INDEX
Straight-ahead linacs, beam design, 115, See also Bent-beam linacs Support assembly, See Patient support assembly Symbols, 298-299 Symmetrical instabilities in field flatness, 24 Synchronous phase, 262
Target volume, 22 Tees, microwave, 98-99 Terminating radiation, 169 Terminology, 300-308 Test equipment/instrumentation,maintenance of, 254-255 Tissue compensators, 205-207 Total-bodyfiemibody x-ray therapy, 35-37 Total skin electron therapy, 38-39 Transformer-rectifier units, 268-270 Transit time, 64-66 Transmission lines, 5 1-53 types of, 52,53 Traveling electromagnetic wave, 9 Traveling-wave accelerators, 70-76 beam loading and load line, 75-76 electron injection and bunching, 72-75 operation, theory of, 70-7 1 standing-wave accelerators compared to, 82-86 structures, 7 1 Treatment, radiotherapy modalities, 33-48 Treatment accessories, 204-209 applicator, electron, 37,40, 143- 144 field shaping systems, 204-205 mechanical/optical pointers, 208 patient contour systems, 207 patient immobilization devices, 208 patient positionlmotion detection, 208-209 wedge filtersltissue compensators, 138,205-207 Treatment beam generation, 4-6, See also Treatment beam production electron beams, 2,778-279 radiation interactions in patient, 279-281 x-ray beams, 275-278 Treatment beam production, 137- 156, See also Treatment beam generation ancillary components, 139- 142 beam collimators, 141, See also Collimation, collimator, multileaf collimator field light and rangefinder, 142 radiation shielding, 140- 141 electron therapy, 142- 145 beam current requirements, 163 beam requirements and subsystem, 142- 143 electron scanning system, 43, 144- 145 electron scattering system, 144 microtrons vs. linacs for, 145
neutron leakagelradioactivation,150- 151 radiation beam, contamination of, 150 radiation head, geometric restrictions of, 138-139 bent-beam vs. straight-thru design, 139 scanned beam dosimetry, 150 treatment head, See radiation head x-ray therapy, 146- 150 beam characteristics and subsystem, 146 x-ray scanning system, 43, 148-150 x-ray target and flattening filter, 147- 148 Treatment beam stabilization, 164-166, See also Dose monitoring requirements, 164 Treatment chair, 201,203-204 Treatment course fractionation, machine performance requirements and, 2-3 Treatment field symmetry, 12 Treatment modalities, See Radiotherapy modalities Treatment planning, 220-224 computer and, 222 definition of, 220 dose contributions, computation of, 220-221 resources, 222-224 simulator verification and, 221 Treatment prescription, 157,221-222 Treatment room design, megavoltage therapy accelerator facilities, 247 Treatment simulation, 10 Treatment simulators, 213-220,257 contemporary developments, 219-220 mechanical features, 2 14-2 16 operational organization, 2 17 radiography/fluoroscopy, 216 regulatory requirements, 2 17-2 19 simulation accessories, 2 17 use of, 2 19 Treatment tablelcouch: top, 203 types of, 201-203 Triode guns, 191 Tumor localization, 10 Tumor volume, 22
Units, 299-300
Vacuum systems, 107- 110 Van de Graaff generator, 270-271 Verification radiograph, 224 Voltage breakdown, 269,271 Voltage standing-wave ratio, 55
Water cooling system, 110- 111 Water loads, microwave, 101- 102
315
316
Waveguides: bends and twists, 97-98 flexible, 97-98 See also Accelerator structures Waveguide windows, 100- 101 Wedge filtersltissue compensators, 138,205-207
X-ray beam generation, 275-278, See also 146-150 angular distribution of photon intensity, 275-276 photon spectra on axis of unflattened lobe, 275 target material/thickness:
INDEX
choice of, 276-278 x-ray scattering, 278 X-ray energy, 3-4 specification, 140, 146 X-ray therapy, 146- 150 beam current requirements in, 18, 163 megavoltage, 34-35 orthovoltage, 33-34 scanned beam dosimetry, 150 scanning system, 148- 150 total-body/hemibody, 35-37 x-ray target and flattening filter, 147-148
